PURASORB PLDL 8038 Medical Device L/DL-Lactide Copolymer is supplied as an amorphous copolymer of L-lactide and DL-lactide with a molar ratio of 70/30 and an inherent viscosity midpoint of 3.8 dL/g when measured as a 0.1% w/v solution in chloroform at 25°C according to ISO 1628-1:2009. The polymer is manufactured under current good manufacturing practice for medical device raw materials and is intended for resorbable implant and drug delivery applications in which bulk hydrolytic degradation follows a period of initial mechanical function. The amorphous character derives from the DL-lactide units, which disrupt stereoregularity and suppress crystallization; the material therefore exhibits no melting endotherm and has a glass transition temperature in the range of 55°C to 60°C, a property that controls dimensional stability at body temperature and influences cold-crystallization avoidance during molding.

Because the copolymer is hygroscopic and hydrolytically sensitive, the product is delivered as white to off-white cylindrical granules in sealed multilayer aluminum foil pouches under vacuum or inert gas. Residual lactide monomer is controlled to ≤ 0.5 wt%, residual tin from the stannous octoate catalyst is specified below 100 ppm, and moisture content is typically kept below 0.5 wt% on the certificate of analysis. Granules exposed to ambient air at relative humidity above 60% require pre-drying before melt processing; recommended drying conditions are 80°C to 100°C for 4 h to 8 h in a vacuum oven or dehumidifying hopper with a dew point below -40°C.

What distinguishes PLDL 8038 from lower-viscosity 70/30 L/DL-lactide copolymers?

The primary distinction is molecular weight expressed as inherent viscosity. A lower-viscosity grade such as a 2.8 dL/g PLDL has shorter polymer chains, which produce lower melt elasticity, faster melt flow, and more rapid moisture-induced chain scission during storage and processing. PLDL 8038 at 3.8 dL/g provides a longer degradation timeframe and higher initial mechanical strength in amorphous matrices, but the processing window narrows because the melt viscosity under ISO 1133-1:2022 conditions is substantially higher. For injection molding, the higher molecular weight requires greater plastication torque and longer screw recovery, while for solution spinning or electrospinning it reduces the maximum solution concentration that can be processed at a given target viscosity. The grade suffix therefore encodes the viscosity midpoint rather than the composition; composition is stated on the certificate of analysis as the L/DL molar ratio.

The values in the accompanying table are representative of the medical device grade; the current certificate of analysis governs individual lot acceptance.

Batch release of PLDL 8038 normally includes determination of inherent viscosity in chloroform, residual monomer by gas chromatography, residual tin by inductively coupled plasma optical emission spectrometry, and moisture by Karl Fischer titration. The glass transition temperature is not usually a pass-fail release specification because it is sensitive to thermal history and moisture; it is reported as typical information. Molecular weight parameters determined by size-exclusion chromatography may be reported on the certificate of analysis when requested. Because the polymer is amorphous, the melt flow index is not a standard release test; inherent viscosity is the primary rheological indicator used for lot-to-lot comparison.

When melt processing exceeds 200°C or residence time exceeds 15 min

At melt barrel set points above 200°C, the rate of chain scission and lactide reformation increases measurably, and the polymer can become tacky and discolored. The recommended melt-processing window for PLDL 8038 is therefore 170°C to 200°C, with the upper limit constrained by thermal degradation rather than by a melting transition. Because the polymer is amorphous, no melting heat is required; the material softens across the glass transition and flows under shear. In a 25 mm co-rotating twin-screw extruder with an L/D ratio of 40:1, a flat temperature profile of 180°C and screw speed of 200 rpm typically produces melt pressure at the die in the range of 3 MPa to 5 MPa. Raising screw speed above 250 rpm can cause local melt temperature overshoot above 210°C, which accelerates molecular weight loss and increases residual lactide. Therefore, the residence time is maintained below 10 min where possible, and the melt temperature is monitored with an exposed-tip thermocouple rather than inferred from barrel set points.

Drying is critical because water acts as a nucleophile and hydrolyzes ester bonds during plastication. A residual moisture content above 0.05 wt% can reduce molecular weight enough to shift inherent viscosity by more than 0.1 dL/g during a single melt pass. Vacuum drying at 80°C for 8 h is common, but the granules should be spread in trays no deeper than 3 cm to allow uniform moisture diffusion. Desiccant or dehumidifying hopper drying at 80°C with a dew point below -40°C is recommended for continuous molding. Overnight drying above 100°C should be avoided because thermal degradation can occur even in the solid state near the glass transition.

On a hydraulic injection molding machine with 50 t clamp force and 25 mm screw diameter, molding PLDL 8038 into a 3.2 mm thick tensile bar typically uses barrel zones of 175°C, 185°C, and 195°C from feed to nozzle, back pressure of 0.5 MPa to 1.0 MPa, and screw rotation below 80 rpm to limit shear heating. Batch-to-batch variation in inherent viscosity from 3.6 dL/g to 4.0 dL/g changes melt cushion stability; the nozzle zone is therefore held to ±5°C when processing near the upper molecular weight limit. Mold temperature is kept low, typically 25°C to 45°C, to reduce cycle time, but too-low mold temperature can freeze orientation stresses into the part and increase warpage risk. In injection molding, screw design is a moderate-compression, low-shear geometry with a compression ratio of 2.0:1 to 2.5:1 and a non-return ring clearance checked against the manufacturer’s tolerance. High compression ratios and intensive mixing sections can generate excessive shear heating in a high-viscosity grade such as PLDL 8038. The shot size should be kept between 30% and 70% of barrel capacity to avoid long residence time in the melt. If machine stoppage exceeds 10 min, the barrel should be purged with a lower-viscosity purge compound or fresh polymer; the melt should not be held at full temperature during unplanned downtime because molecular weight loss continues in the absence of shear.

Sterilization pathways and molecular weight retention

For gamma sterilization at 25 kGy to 40 kGy according to ISO 11137-1:2006, the amorphous PLDL 8038 undergoes dose-dependent chain scission that is reflected in a drop in inherent viscosity. The magnitude depends on initial moisture content, packaging atmosphere, and irradiation temperature; published data for this specific configuration is limited, and final device testing is required to establish an acceptable dose range. In inert packaging at low moisture, the viscosity loss is typically lower than in air, but the loss is not zero. Ethylene oxide sterilization at 45°C to 55°C is less hydrolytically aggressive than steam, but it requires extended aeration to meet residual ethylene oxide limits under ISO 10993-7:2008. Steam sterilization is contraindicated because the combination of heat and water initiates bulk hydrolysis within the polymer. Sterilization compatibility should be evaluated on the final packaged device, and the molecular weight after aging should be compared with the mechanical requirements of the implant.

Hydrolytic degradation in phosphate-buffered saline at pH 7.4 and 37°C proceeds by bulk erosion rather than surface erosion for this copolymer. Water uptake precedes ester bond cleavage, and the amorphous matrix loses mass after a molecular weight threshold is reached. In vitro mass loss for a 70/30 L/DL-lactide copolymer with an inherent viscosity midpoint of 3.8 dL/g typically becomes measurable after 8 to 12 weeks, while complete mass loss may extend beyond 12 months depending on part cross-section, residual monomer, and sterilization history. However, published data for this specific configuration is limited; degradation studies must be repeated on the final sterilized device geometry because gamma irradiation can shorten the mass-loss onset. The bulk erosion mechanism creates a lag phase during which mechanical properties may decline before mass loss is detectable. Tensile strength retention of amorphous 70/30 PLDL at 37°C in buffered saline is typically several weeks to a few months, but the exact duration depends on the initial number-average molecular weight, the degree of orientation, and the presence of acidic degradation products. Because the copolymer degrades into lactic acid, the local pH within a large dense implant can fall, potentially accelerating autocatalytic hydrolysis in the core. This autocatalytic effect is reduced in porous or thin-walled devices where degradation products can diffuse out. Manufacturers should monitor pH, inherent viscosity, and mass loss simultaneously when developing resorbable implants.

When evaluated against semi-crystalline poly(L-lactide), PLDL 8038 differs in morphology, thermal behavior, and degradation profile. The DL-lactide co-units suppress crystallization, so the material does not form the spherulitic structure responsible for anisotropic shrinkage and late-stage crystalline micro-debris observed with PLLA. The absence of crystallinity also reduces the tensile modulus relative to PLLA and eliminates the fatigue crack propagation associated with crystalline/amorphous boundaries. Tensile properties of compression-molded amorphous specimens measured according to ISO 527-2:2012 typically fall in the range of 45 MPa to 60 MPa for tensile strength and 2.5 GPa to 3.0 GPa for tensile modulus, but these values depend on molecular weight, moisture, and specimen preparation. Elongation at break in the amorphous state is generally below 10% for unoriented specimens, making the material stiff and brittle compared with polycaprolactone. Published data for this specific configuration is limited; mechanical testing on final devices is required. Against faster-degrading poly(lactide-co-glycolide) 50/50, PLDL 8038 retains mechanical function longer and releases acidic degradation products more slowly, which can reduce local pH drop in confined implant sites. Compared with a lower-viscosity 70/30 PLDL at 2.8 dL/g, PLDL 8038 can provide higher initial tensile strength and longer strength retention, but the lower-viscosity grade is easier to injection mold into thin-wall or complex geometries. The choice between the two is therefore a trade-off between mechanical lifetime and processability. Compared with poly(L-lactide) homopolymer of the same viscosity class, PLDL 8038 has lower modulus and no crystalline reinforcement, but it avoids the long-term crystalline particle release associated with PLLA.

Compliance documentation for PLDL 8038 is generated under a quality system certified to ISO 13485:2016. The polymer is supplied with data to support biocompatibility evaluation under ISO 10993-1:2018, including cytotoxicity, sensitization, irritation, acute systemic toxicity, genotoxicity, implantation, and subchronic systemic toxicity endpoints. Device manufacturers may reference the polymer documentation in submissions under FDA 21 CFR Part 820 and EU Medical Device Regulation 2017/745, but the responsibility for final device validation remains with the legal manufacturer. The polymer is not supplied as a finished device and is not certified for direct clinical use without additional processing and evaluation.

For interference screws and fracture fixation pins, PLDL 8038 is injection molded into amorphous devices that rely on oriented molecular chains to achieve higher mechanical strength. The mold design typically includes a gate size sufficient to maintain packing pressure during cooling; premature gate freeze-off can reduce part density and shift mechanical properties. Because the material is amorphous, mold shrinkage is lower and more isotropic than PLLA, but shrinkage values of 0.3% to 0.8% are still accounted for in tool dimensions. In soft-tissue suture anchors, the polymer is injection molded or machined from molded preforms. Machining produces sharp edges and localized heating; cutting tools must be sharp and speeds controlled to avoid melting or stress whitening. A machined PLDL 8038 anchor is typically inspected for burrs and microcracks because the amorphous glassy state is notch-sensitive. Mechanical pull-out strength depends on thread geometry and insertion technique rather than polymer properties alone, but the polymer must maintain strength through the insertion torque and the early healing period. Published data for this specific configuration is limited.

In a drug-eluting implant matrix, PLDL 8038 is solution-processed in chloroform or dichloromethane at polymer concentrations between 2% w/v and 10% w/v, followed by casting, spray drying, or solvent extrusion. Release of a hydrophobic active pharmaceutical ingredient from the amorphous matrix is controlled by diffusion through the polymer followed by matrix erosion; initial burst release is influenced by drug loading, drug solubility in the polymer, and solvent removal rate. Mechanical properties after drug loading are measured according to ISO 527-2:2012 or ASTM D638-14, and the degradation profile is re-established on the sterilized drug-loaded device because the active ingredient and sterilization step can alter glass transition, moisture uptake, and chain-scission kinetics.