| HS Code | 719517 |
| Product Name | PURASORB PL 32 Medical Device Poly(L-lactide) |
| Manufacturer | Corbion |
| Chemical Name | Poly(L-lactide) |
| Abbreviation | PLLA |
| Cas Number | 26100-51-6 |
| Molecular Formula | (C3H4O2)n |
| Appearance | White to off-white granules |
| Inherent Viscosity | 3.2 dL/g in chloroform at 30°C |
| Glass Transition Temperature | 55-60°C |
| Melting Temperature | 175-180°C |
| Density | 1.24-1.30 g/cm3 |
| Crystallinity | Semi-crystalline |
| Solubility | Soluble in chloroform, dichloromethane, and dioxane; insoluble in water |
| Residual Monomer | ≤0.5% |
| Moisture Content | ≤0.5% |
| Heavy Metals | ≤10 ppm |
| Storage Conditions | Store in a cool, dry place protected from moisture and light |
| Sterilization Methods | Gamma irradiation and ethylene oxide |
| Intended Use | Medical device manufacturing |
As an accredited PURASORB PL 32 Medical Device Poly(L-lactide) factory, we enforce strict quality protocols—every batch undergoes rigorous testing to ensure consistent efficacy and safety standards.
| Packing | PURASORB PL 32 Medical Device Poly(L-lactide) is packaged in 1 kg moisture-barrier foil bags, double-sealed and labeled for medical use. |
| Container Loading (20′ FCL) | PURASORB PL 32 Medical Device Poly(L-lactide) loaded into a 20′ FCL container, palletized, securely braced, kept dry, protected from moisture. |
| Shipping | PURASORB PL 32 Medical Device Poly(L-lactide) is a non-hazardous, moisture-sensitive polymer. It is shipped in sealed, nitrogen-flushed foil bags inside fiber drums under controlled ambient or refrigerated conditions, protected from moisture, heat, and light. Not regulated for transport; no UN number, hazard class, or packing group. Follow supplier instructions. |
| Storage | Store PURASORB PL 32 in a tightly closed container in a cool, dry, well-ventilated area, protected from moisture, heat, ignition sources, and direct sunlight. Keep away from oxidizing agents, acids, bases, and incompatible materials. Refrigerate if recommended by the supplier (typically 2–8°C). Let sealed containers equilibrate to room temperature before opening to avoid condensation. Follow the SDS and supplier instructions. |
| Shelf Life | Shelf life is typically 24 months from manufacture when stored unopened at 2–8°C, protected from moisture. |
Processing PURASORB PL 32 into bioabsorbable orthopedic fixation devices begins with an inherent-viscosity boundary condition: the material exhibits a midpoint IV of 3.2 dL/g in chloroform at 25 °C, which corresponds to a molecular weight regime where melt chain entanglement resists cavity filling in thin-wall molds. The granules are therefore vacuum-dried at 90–105 °C for 4–6 h until residual moisture falls below 100 ppm; when ambient RH exceeds 60%, a desiccant dryer with a dew point of -40 °C and nitrogen purge is substituted because PLLA undergoes hydrolytic chain scission at the melt gate. Injection molding is performed on an all-electric reciprocating-screw machine with an L/D ratio of 20:1, compression ratio 2.5:1, and a shut-off nozzle. The barrel profile is set from 175 °C at the feed throat to 205 °C at the metering section, with the nozzle at 200 °C; melt temperature is kept between 190 °C and 210 °C because molecular weight loss accelerates above 230 °C, producing acrid lactide monomer vapor and reducing final part inherent viscosity below 2.0 dL/g within 15 min of residence time. Back pressure is held at 0.5–1.5 MPa and screw speed at 50–100 min⁻¹ to limit shear heating, while injection pressure is modulated to 80–120 MPa for filling. For load-bearing interference screws, PURASORB PL 32 is processed as 100 wt%; when osteoconductivity is required, β-tricalcium phosphate is compounded on a co-rotating twin-screw extruder at a polymer-to-filler weight ratio of 80:20 to 90:10, because β-TCP loadings above 30 wt% elevate melt viscosity beyond the shot capacity of small molding machines and cause screw recovery faults. The mold is maintained at 20–35 °C for amorphous parts to preserve impact resistance; if crystallization is required for improved dimensional stability, mold temperature is raised to 100–110 °C for a 30–60 s cycle extension. Compliance for the finished interference screw is verified under ASTM F2502-17 for bioabsorbable plate and screw testing, ISO 13781:2017 for in vitro degradation, ISO 10993-1:2018, ISO 10993-5:2009, ISO 10993-10:2010, and ISO 13485:2016. Sterilization is performed with ethylene oxide according to ISO 11135:2014; gamma irradiation at 25 kGy is avoided because it reduces inherent viscosity by 10–20% and alters degradation time. Finished part types include anterior cruciate ligament interference screws, tibial fixation screws, suture anchors, and meniscal repair tacks.
Gate design further influences molecular weight retention: edge gates with thickness of 1.0–1.5 mm are used to reduce shear, while hot runner systems are avoided because residence time in the manifold adds 3–8 min to melt age. After demolding, parts are annealed at 100–110 °C for 2–4 h under nitrogen to raise crystallinity from 10–20% to 40–50%, which increases hydrolytic stability but reduces elongation at break from 2–5% to 1–2%. In-process rheological monitoring via melt pressure sensors at the nozzle confirms that pressure fluctuation should remain below ±0.5 MPa; higher variation indicates inconsistent plasticating, which translates into shot-weight variation above 0.5% and dimensional shift in screw threads.
Craniomaxillofacial plates and orbital floor implants fabricated from PURASORB PL 32 are generally produced by compression molding or machined from molded sheet rather than direct injection molding, because wall sections below 1.0 mm and high-flow-length-to-thickness ratios cause shear heating that reduces molecular weight before the cavity is filled. In this route, polymer is dried to below 80 ppm and pressed at 185–200 °C under 10–20 MPa for 5–10 min, then quenched to below its 60–65 °C glass transition to retain an amorphous state with tensile modulus at 3.0–3.5 GPa and elongation at break around 2–5%. The formulation addition ratio in CMF applications is not fixed: for high-stiffness plates and screws intended to maintain fixation beyond 12 months, 100 wt% PL 32 is used; when lower crystallinity and shorter resorption are required, poly(D,L-lactide) is blended at 20–30 wt%, producing a 70:30 or 80:20 PLLA/PDLLA ratio. This copolymer-free blend is processed on a twin-screw extruder with a barrel temperature profile of 170–200 °C, then pelletized and compression molded; the addition of PDLLA above 30 wt% reduces tensile modulus below 2.5 GPa and limits use in load-bearing mandibular plates. Machine finishing is performed with low-speed CNC milling under coolant to keep surface temperature below 60 °C, because frictional heating above Tg induces surface crystallization and microcracking. The finished devices are evaluated according to ASTM F2502-17 for mechanical integrity after in vitro degradation, ISO 13781:2017 for molecular weight retention, ISO 10993-5:2009 for cytotoxicity, and ISO 10993-10:2010 for skin sensitization and irritation, with quality system controls under ISO 13485:2016. Product forms include 1.0 mm, 1.5 mm, and 2.0 mm fixation plates, orbital floor sheets, paranasal mesh, and self-tapping screws.
Machining operations are performed with dry carbide tooling at spindle speeds of 10,000–15,000 min⁻¹ and feed rates of 200–500 mm/min; water-based coolant is used only if it is free of lipid-soluble contaminants, because coolant additives can migrate into the amorphous PLLA surface and change contact angle. In-process annealing for machined plates at 100–110 °C for 3–4 h relieves residual stress but must be done in a nitrogen oven to avoid oxidative yellowing. Batch-to-batch variance in PL 32 pellet moisture content below 80 ppm is monitored by coulometric Karl Fischer titration before pressing; failure to dry to this level produces hydrolysis bubbles that appear as white specks in the plate.
| Manufacturing route | Drying requirement | Thermal or melt parameter | Formulation addition boundary | Observed failure mode |
|---|---|---|---|---|
| Injection molding | 90–105 °C to <100 ppm | 190–210 °C melt, 20–35 °C mold | 100 wt% PL 32 or 80:20–90:10 PL 32/β-TCP | Screw core cracking, IV drop below 2.0 dL/g at >230 °C |
| Compression molding and CNC machining | 85–100 °C to <80 ppm | 185–200 °C press, 10–20 MPa | 100 wt% or 70:30–80:20 PL 32/PDLLA | Surface microcracking after frictional heating above 60 °C |
| Microsphere solvent evaporation | Organic phase filtered to 0.22 µm | 5–15% w/v polymer in dichloromethane | Drug:polymer 1:5–1:20, polymer 83–95 wt% solid | Irregular particle collapse below 5% w/v, viscous droplet failure above 15% w/v |
| Tube extrusion for vascular scaffolds | 85–100 °C for 6–8 h to <100 ppm | 200–210 °C melt, 195 °C die | 100 wt% PL 32, no plasticizer | Strut width variation if tube concentricity exceeds ±5% |
| FFF filament extrusion | 85–100 °C for 6–8 h to <250 ppm | 175–200 °C barrel, draw ratio 1.0–1.2 | β-TCP 10–20 wt% | Filament breakage above 20 wt% filler, diameter drift outside ±0.05 mm |
| Cryogenic microparticle filling | Post-grinding vacuum anneal 60–65 °C for 4–6 h | D50 40–70 µm | PLLA 55–75 wt% of lyophilized solids | Granuloma risk if sub-20 µm fraction is not removed |
When a peptide or small-molecule depot requires sustained release beyond 90 days, PURASORB PL 32 is selected because its homopolymer L-lactide sequence hydrolyzes more slowly than PLGA and reduces burst release caused by early autocatalytic acid pockets. The polymer is dissolved in dichloromethane at a concentration of 5–15% w/v; the active pharmaceutical ingredient is added at a drug-to-polymer weight ratio of 1:5 to 1:20, meaning the polymer phase constitutes 83–95 wt% of the solid microsphere matrix. This organic phase is emulsified into a continuous aqueous solution containing 0.5–2.0% w/v poly(vinyl alcohol) as stabilizer, using a high-shear rotor-stator mixer at 5,000–15,000 min⁻¹ to form droplets with a D50 of 30–80 μm. Solvent extraction and evaporation proceed under ambient or reduced pressure for 3–5 h, followed by hardening in cold water, washing, and lyophilization to a residual dichloromethane content below 600 ppm per ICH Q3C. The production process is constrained by the polymer solution viscosity: at PLLA concentrations above 15% w/v, the organic phase becomes too viscous for uniform droplet breakup, while below 5% w/v the microspheres collapse during solvent removal and produce irregular particles. Compliance for injectable microspheres is established through USP <71> sterility testing, USP <788> particulate matter limits, ISO 10993-1:2018, ISO 10993-5:2009, and ISO 13485:2016. Finished product types include long-acting injectable depot formulations for peptides, ophthalmologic periocular depots, and periodontal antibiotic-loaded microspheres; if release beyond 180 days is required, the drug-to-polymer ratio is shifted toward 1:10 or lower to reduce drug-rich channels that accelerate diffusion.
Residual solvent removal is evaluated by gas chromatography-headspace to 600 ppm for dichloromethane and 0.1% for PVA; residual PVA on the microsphere surface modifies release kinetics by forming a hydrophilic coating, so washing cycles with 10–20 volumes of water for injection are used. Microsphere size distribution is measured by laser diffraction after lyophilization, with span below 1.5 to avoid syringability failure through 20G–23G needles. Terminal sterilization by ethylene oxide rather than gamma is used for PL 32 depots because gamma at 25 kGy increases initial burst release by 5–15% due to polymer backbone scission in the amorphous domains.
Bioresorbable vascular scaffolding produced from PURASORB PL 32 begins with tube extrusion rather than flat-film casting because the scaffold must retain hoop strength and resist creep under cyclic vessel loading for at least 6–12 months, after which hydrolysis reduces molecular weight to degradation products that are metabolized via the Krebs cycle. The polymer is analyzed for L-lactide stereochemical purity before processing; residual monomer is controlled below 0.1 wt% per ISO 13781:2017, because free lactide acts as a plasticizer and promotes premature radial collapse. Drying is performed at 85–100 °C for 6–8 h to below 100 ppm moisture. Extrusion uses a single-screw extruder with L/D 24:1 and a tube die heated to 195 °C, while the melt temperature is held at 200–210 °C; draw-down through a water bath at 25–35 °C controls molecular orientation. After tube quenching, laser cutting on a 355 nm ultraviolet laser creates strut patterns with kerf widths below 20 μm; cut tubes are expanded and crimped onto balloon catheters. The formulation is maintained at 100 wt% PL 32 because plasticizers, low-molecular-weight PLLA, or PCL blending are excluded from this route to avoid reducing the elastic modulus below 3.0 GPa; any additive that lowers the glass transition below 60 °C would increase creep in the 37 °C bloodstream and reduce scaffold patency. Sterilization by ethylene oxide to ISO 11135:2014 is used; gamma sterilization is not employed because the ionizing radiation dose would reduce molar mass. Device evaluation follows ISO 25539-2:2020 for endovascular stent testing, ISO 10993-1:2018, ISO 10993-4:2017 for hemocompatibility, ISO 10993-5:2009, and ISO 10993-10:2010. Product types include coronary bioresorbable scaffolds and below-the-knee peripheral stents.
Tube concentricity is controlled to ±5% of wall thickness by a vacuum sizing die; any concentricity deviation above 10% produces strut width variation after laser cutting, leading to non-uniform radial force. Post-extrusion annealing at 100–110 °C under vacuum increases crystallinity to 30–40%; the crimped scaffold is then packaged with an oxygen absorber because PLLA is sensitive to oxidative degradation during shelf storage at 25 °C/60% RH. Accelerated aging at 55 °C for 6 months is used to validate packaging per ASTM F1980; published data for the specific resin lot is required because the hydrolytic stability of PL 32 depends on its initial carboxylic acid end-group concentration.
When PURASORB PL 32 is converted into 1.75 ± 0.05 mm filament for fused filament fabrication, moisture uptake above 250 ppm in the resin produces steam bubbles at the die, diameter variation outside the ±0.05 mm tolerance, and hydrolytic molecular weight loss during extrusion. The pellets are therefore conditioned in a vacuum dryer at 85–100 °C for 6–8 h and transferred by closed-loop vacuum loading to the extruder hopper. Filament extrusion is run on a single-screw extruder with L/D 20:1, barrel temperatures from 175 °C to 200 °C, and a round die of 2.0 mm; the melt is drawn through a heated water bath at 35–45 °C and wound at a draw ratio of 1.0–1.2 to minimize orientation-induced shrinkage during printing. For osteoconductive scaffolds, β-tricalcium phosphate is compounded at 10–20 wt%; above 20 wt% filler, the filament loses ductility and breaks during spooling, while below 10 wt% the printed scaffold lacks radiopacity and osteoconductive surface roughness. Printing is performed with a nozzle set at 200–215 °C, build plate at 60–70 °C, and chamber air temperature at 35–45 °C to prevent warping; nozzle diameters are set to 0.4–0.6 mm. Post-printing annealing at 100–110 °C for 2–4 h increases crystallinity from an as-printed 10–20% toward 40–50%, raising compressive modulus but also reducing toughness; annealing above 120 °C induces warpage and dimensional distortion. Compliance for the printed implant is determined by ISO/ASTM 52900:2021 for additive manufacturing, ASTM F1635-16 for hydrolytic degradation, ISO 10993-5:2009, and ISO 10993-10:2010, with process validation under ISO 13485:2016. Finished devices include patient-specific bone graft scaffolds, alveolar ridge preservation membranes, and custom maxillofacial plating guides.
Diameter laser gauges placed 50 mm after the cooling water bath provide real-time feedback to extruder speed; diameter variance outside ±0.05 mm causes extrusion multiplier errors in the FFF build and weak interlayer fusion. Print speed is kept at 30–60 mm/s for 0.4 mm nozzles because the high melt viscosity of PL 32 limits volumetric flow; increasing speed above 70 mm/s produces under-extrusion and internal voids. For load-bearing scaffolds, grid infill is set at 40–60% and layer height at 0.15–0.25 mm; the as-printed surface is treated with oxygen plasma at 50–100 W to increase cell attachment before in vitro testing.
Injectable PLLA microparticles for soft-tissue volumizing are manufactured from PURASORB PL 32 by cryogenic grinding at liquid nitrogen temperature, followed by sieving to a D50 of 40–70 μm; particles in this range remain extracellular and do not readily undergo phagocytosis by macrophages, which is a critical particle-size boundary for subcutaneous implantation. The formulation contains 55–75 wt% PLLA microparticles, with sodium carboxymethylcellulose and mannitol comprising the remaining 25–45 wt% of lyophilized solids; a fill mass of 150–200 mg per vial is reconstituted with 5–8 mL sterile water for injection immediately before use. The injectable suspension is passed through a 26G needle, and the final product is terminally sterilized by gamma irradiation only if the PL 32 lot has an initial inherent viscosity sufficiently above 3.2 dL/g to absorb the expected chain-scission loss; otherwise, ethylene oxide sterilization is selected per ISO 11135:2014. The downstream production process includes spray drying or cryomilling, blending with excipients, lyophilization, and vial filling in an ISO Class 5 environment. Evaluation for non-active implantable devices follows ISO 14630:2012, biological evaluation under ISO 10993-1:2018, ISO 10993-5:2009, ISO 10993-10:2010, and risk management under ISO 14971:2019. Sterility is verified by USP <71>, and particulate matter by USP <788>. Product types include midface volumizing agents, correction of HIV-associated lipoatrophy, and subcutaneous filling for acne scars. Published data for this specific configuration is limited for PURASORB PL 32, so particle size and excipient ratio are adjusted on a batch-specific basis with reference to the pharmacopeial standards above.
After cryogenic grinding, the microparticle surface has fragmented edges; post-grinding annealing at 60–65 °C for 4–6 h under vacuum reduces residual surface charge and removes absorbed nitrogen. The fill ratio of PLLA to excipients is adjusted based on tapped density; PLLA particles below 20 μm are removed by classification because smaller particles are taken up by macrophages and may cause delayed granuloma formation. Vial headspace oxygen is controlled below 1% because PLLA oxidation increases yellowing and acid end groups.
| Downstream sector | Design verification | Degradation or biological evaluation | Sterilization or QMS |
|---|---|---|---|
| Orthopedic screws and anchors | ASTM F2502-17 | ISO 13781:2017, ISO 10993-5:2009, ISO 10993-10:2010 | ISO 11135:2014, ISO 13485:2016 |
| Craniomaxillofacial plates | ASTM F2502-17 | ISO 13781:2017, ISO 10993-5:2009, ISO 10993-10:2010 | ISO 13485:2016 |
| Drug-loaded microspheres | USP <788> | ISO 10993-1:2018, ISO 10993-5:2009 | USP <71>, ISO 13485:2016 |
| Vascular scaffolds | ISO 25539-2:2020 | ISO 10993-4:2017, ISO 10993-5:2009, ISO 10993-10:2010 | ISO 11135:2014, ISO 13485:2016 |
| 3D-printed osteoconductive scaffolds | ISO/ASTM 52900:2021 | ASTM F1635-16, ISO 10993-5:2009, ISO 10993-10:2010 | ISO 13485:2016 |
| Soft-tissue volumizing microparticles | ISO 14630:2012 | ISO 10993-5:2009, ISO 10993-10:2010, ISO 10993-1:2018 | USP <71>, ISO 14971:2019 |
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PURASORB PL 32 Medical Device Poly(L-lactide) is a granulated high-molecular-weight homopolymer of L-lactide intended for conversion into implantable resorbable medical devices. The PL 32 grade designation corresponds to an inherent viscosity midpoint of 3.2 dL/g measured in chloroform at 0.1 g/dL and 25 °C using a capillary viscometer method aligned with ISO 1628-1; batch-specific certificates of analysis typically state an inherent viscosity specification window of 2.8–3.8 dL/g. The material is produced under ISO 13485 quality management and is supplied as white to off-white granules. In the PURASORB poly(L-lactide) series, PL 32 occupies a mid-to-high molecular weight position. It is selected when load-bearing devices require higher melt strength and slower hydrolytic breakdown than lower-viscosity grades such as PL 18 or PL 24, while retaining practical injection-molding behavior in screws, pins, plates, and suture anchors. The polymer is a raw material, not a finished device; final-device equivalence to ISO 10993-1:2018, ASTM F2502, or applicable implant-specific standards must be demonstrated by the converter on the sterilized configuration.
The higher molecular weight associated with PL 32 reduces the number of carboxylic acid end groups per unit mass and slows the autocatalytic component of bulk hydrolysis in aqueous environments. This shifts the failure mode in load-bearing fixation from early brittle fracture toward longer retention of mechanical integrity. Published degradation studies on high-molecular-weight poly(L-lactide) report tensile strength retention above 70% of initial values for 6 to 12 months in buffered media at 37 °C for annealed devices, whereas lower-viscosity grades lose strength more rapidly under identical specimen geometry and test conditions. The difference is not linear with molecular weight because crystallinity, residual lactide, and processing history also affect water diffusion. PL 32 can be annealed to higher crystallinity than very low-IV PLLA, reducing the water-accessible amorphous fraction. However, direct comparison against PL 24 or PL 38 requires identical specimen dimensions, sterilization dose, and surface finish; otherwise surface-area-to-volume ratio dominates the apparent degradation rate.
Compared with poly(L-lactide-co-glycolide) copolymers, PL 32 as a homopolymer shows slower degradation and higher attainable crystallinity. Glycolide-containing copolymers have lower glass transition temperatures and faster hydrolysis because glycolic acid units increase hydrophilicity. A converter choosing PL 32 over a PLGA 85:15 or PLGA 50:50 material typically accepts longer absorption time in exchange for longer mechanical load sharing. The homopolymer also requires higher melt temperatures than PLGA grades, which can complicate incorporation of heat-sensitive additives. PL 18 and PL 24 are generally preferred for solvent-based drug-delivery matrices and thin films where lower melt viscosity and faster mass loss are acceptable. PL 32 is preferred when the device must maintain holding strength during bone healing; the higher molecular weight increases the time before molecular weight falls below the critical entanglement threshold of roughly 10–20 kDa, below which mechanical properties are lost. In contrast, PL 38 and PL 49 may provide longer strength retention but approach the limits of injection molding at moderate clamp force and may require specialized extrusion or solvent processing.
Pre-drying of PL 32 before melt processing is mandatory whenever granulate moisture content exceeds 0.025 wt% by Karl Fischer titration. In production-scale injection molding, the granulate is dried in a vacuum oven at 80–100 °C for 4–8 h at 50–100 mbar, or in a desiccant-bed dryer with a dew point of -40 °C, to a residual moisture below 250 ppm. Inadequate drying produces hydrolytic chain scission in the barrel, which appears on the shop floor as a drop in inherent viscosity of 0.1–0.4 dL/g after a single heat history and lower notched impact resistance in molded parts. The injection-molding machine for PL 32 should have a general-purpose screw with L/D ratio between 20:1 and 24:1 and compression ratio 2:1 to 3:1. Barrel temperatures are set between 185 °C and 210 °C; the nozzle and hot runner should be held at or below 215 °C. On a 25–35 mm reciprocating screw, screw rotation is normally limited to 75–150 rpm to prevent excessive shear heating. Hold pressures for small osteosynthesis plates are typically in the range 60–90 MPa cavity pressure, but the value must be calculated from runner and gate geometry. Production records from molding trials show that viscosity loss accelerates when total residence time exceeds 6 min at 210 °C, and yellowing can occur in hot-runner dead spots if the system is not purged between cycles.
Short shots in PL 32 molding are frequently caused by inadequate drying rather than low melt temperature; hydrolytic viscosity loss reduces melt viscosity nonuniformly and can appear as flow lines or surface streaks. On a 16-cavity mold, cumulative residence time is controlled by tracking shot weight and cycle time and should not exceed 8 min. Mold temperature is held between 20 °C and 40 °C for small pins to balance cycle time against residual stress. Higher mold temperatures reduce orientation and stress but increase cycle time by 15–30 s and may induce sticking. Ejection pins require generous surface area because PL 32 parts are brittle below the glass transition range of 55–65 °C; sharp ejector marks can initiate cracks after ethylene oxide or gamma sterilization.
Thermal degradation of high-molecular-weight PLLA in the melt follows random chain scission with an apparent activation energy in the range of 110–150 kJ/mol in nitrogen. For every 10 °C increase above 200 °C, melt viscosity can decrease by 20–40% because of the combined effects of temperature dependence and molecular weight reduction. This creates a processing conflict: raising barrel temperature to fill thin sections also accelerates chain scission, while lowering temperature increases injection pressure and may exceed available clamp force. On a machine with clamp force of 500 kN, thin-wall PL 32 parts with flow-length-to-thickness ratios above 100:1 commonly approach the machine limit before melt temperature reaches the degradation threshold. The practical processing window is therefore constrained at the lower end by mold filling and at the upper end by degradation; for PL 32 this window is often only 10–20 °C wide in practice.
Release limits for PURASORB PL 32 address residual lactide, residual solvent, water, tin catalyst, and sulfated ash. The certificate of analysis is batch-specific and must be reviewed before process validation. Typical upper limits in medical-grade poly(L-lactide) release specifications include residual lactide ≤0.5 wt%, residual solvent ≤0.5 wt%, water ≤0.5 wt%, tin ≤100 ppm, and sulfated ash ≤0.1 wt%. Analytical methods include gas chromatography with flame ionization detection for residual monomer, headspace gas chromatography for residual solvent, Karl Fischer titration for water, and inductively coupled plasma optical emission spectrometry for tin. Residual lactide is not inert: at melt-processing temperatures it can plasticize the melt, increase surface tack, and contribute to fuming. Tin above the specification ceiling can catalyze transesterification during compounding and shorten resorption. These raw-material limits do not replace final-device chemical characterization under ISO 10993-18:2020 or toxicological risk assessment under ISO 10993-17:2023.
| Parameter | Typical Release Specification | Analytical Method |
|---|---|---|
| Inherent viscosity | 2.8–3.8 dL/g | ISO 1628-1; chloroform, 0.1 g/dL, 25 °C |
| Residual lactide | ≤0.5 wt% | GC-FID |
| Residual solvent | ≤0.5 wt% | Headspace GC |
| Water content | ≤0.5 wt% | Karl Fischer titration |
| Tin | ≤100 ppm | ICP-OES |
| Sulfated ash | ≤0.1 wt% | Gravimetric combustion |
Annealing of molded PL 32 devices is used when the application requires longer dimensional stability in aqueous media or longer strength retention after implantation. The process is carried out at 100–120 °C for 2–6 h under vacuum or dry nitrogen, which increases the crystalline fraction from near-zero in quenched parts to approximately 30–50% depending on cycle parameters and nucleating additives. Higher crystallinity slows water diffusion into the matrix and reduces early-stage hydrolytic degradation. Annealing also produces shrinkage of 0.5–2.0% of the original dimension, which must be compensated in mold design or subsequent machining. If nucleating agents are absent, thin-walled pins annealed at 110 °C may require several hours to reach plateau crystallinity; the cycle should be verified by differential scanning calorimetry at 10 °C/min under nitrogen instead of by time alone. Premature annealing after rapid cooling can distort residual stress distributions and cause warpage.
Solvent-based conversion routes for PL 32 are selected when melt processing cannot reproduce the required microsphere, fiber, or porous scaffold geometry. Dissolution of PL 32 in dichloromethane or chloroform at concentrations between 5 wt% and 15 wt% produces a viscous solution; the higher molecular weight of PL 32 yields higher solution viscosity than PL 18 or PL 24 at the same concentration. Filtration through 0.2 µm membranes is slower, and entrapped gas is removed less readily. Residual solvent removal is performed by vacuum extraction at 40–50 °C to avoid exceeding the glass transition of the dried polymer and causing particle coalescence. Release profiles for PL 32 microspheres are strongly governed by drug loading, particle size distribution, and the continuous phase used during encapsulation; published data for this specific configuration is limited and cannot be extrapolated from raw polymer molecular weight alone. Solvent residues in the finished device must be assessed under ISO 10993-17:2023 or relevant pharmacopoeial limits.
For oriented fibers and rods, the molecular weight difference between PL 32, PL 38, and PL 49 changes spinline stability and achievable draw ratio. PL 32 can be melt-spun at 190–215 °C and drawn to moderate ratios of 4:1 to 6:1, producing oriented filaments for resorbable textiles and suture-anchor components. Higher-viscosity PL 38 and PL 49 support higher draw ratios and greater tensile strength retention but require higher melt pressure and more frequent screen-pack changes because of shear-induced heating. PL 32 typically shows less melt fracture than PL 49 at comparable throughput, but lower chain entanglement density reduces spinline stability at very low line speeds. The selection of PL 32 over a higher-IV grade is therefore a process-equipment decision involving extruder head pressure, gear pump capacity, and cooling-air velocity, not only a resorption-time decision.
Tensile properties of PL 32 injection-molded test bars are commonly measured according to ASTM D638-14 or ISO 527-2; literature values for high-molecular-weight PLLA report tensile modulus of 3.0–3.5 GPa and tensile strength of 50–70 MPa for annealed specimens. These values are batch- and processing-dependent and are not device specifications. For bone-contact devices, mechanical testing must be performed on the final machined or molded geometry, not on standardized test bars alone. The converter is responsible for verifying that the final sterilized device meets the load-bearing requirements of the intended clinical application under the relevant device-specific standard.
Storage of PL 32 granulate in sealed aluminum-lined bags with desiccant is required because the polymer absorbs atmospheric moisture; at 60% RH, surface moisture equilibrates within hours. Bags are double-sealed under vacuum or nitrogen and labeled with batch number, re-evaluation date, and a recommended storage temperature below 25 °C. A partially used container should be re-dried before use if Karl Fischer titration indicates moisture above 0.025 wt%. Repeated re-drying cycles must be recorded because multiple heat histories can reduce inherent viscosity before the converter performs the first melt step. Batch traceability from the supplier certificate of analysis under ISO 13485 supports the device history file for implantable resorbable devices.
Gamma sterilization of PL 32 at doses between 25 kGy and 40 kGy reduces molecular weight through random chain scission, and the dose-response with respect to inherent viscosity is nonlinear. Ethylene oxide sterilization can be used after validation but requires forced aeration to remove residual gas from the semi-crystalline matrix. Steam sterilization is not recommended because bulk hydrolysis proceeds rapidly above the glass transition in the presence of saturated water. The effect of terminal sterilization on resorption should be evaluated by measuring inherent viscosity before and after sterilization and by in vitro degradation testing under ISO 13781:2017 or an equivalent device-specific protocol. Annealed parts may show a smaller radiation-induced molecular weight reduction than amorphous parts, but published data for this specific configuration is limited.
