Processing of PURASORB PL 24 for orthopaedic interference screws begins with vacuum drying at 80 °C to 100 °C until residual moisture is maintained below 250 ppm; for this high-IV poly(L-lactide) homopolymer, melt-state hydrolysis is the controlling degradation reaction, and failure to reach this moisture threshold lowers molecular weight during plastication before the melt reaches the mould. The dried resin is fed into a reciprocating-screw injection moulding machine with an L/D ratio of 20:1 to 24:1 and a screw compression ratio of 2.5:1 to 3:1; barrel temperatures are typically profiled from 160 °C at the feed throat to 185 °C at the nozzle, with a back pressure of 5 bar to 15 bar and injection velocity set to fill the thread geometry before the flow front freezes. Mould temperature is held at 20 °C to 30 °C to obtain rapid solidification, but this condition produces an amorphous superficial layer over a slow-cooled semi-crystalline core; differential shrinkage between these zones must be offset by gate placement at the screw head rather than at the thread crest. Total residence time in the barrel and hot runner is restricted to 6 min or less because the melt viscosity of 2.4 dL/g PLLA is sensitive to thermal-mechanical chain scission, which appears in the finished screw as a drop in intrinsic viscosity, a reduced molecular weight tail, and lower fatigue resistance. Unplasticized resin is used for load-bearing fixation; if osteoconductive filler is added, β-tricalcium phosphate or hydroxyapatite is compounded at 5 wt% to 20 wt% only after rheological evaluation, because filler raises melt pressure at the nozzle and reduces elongation at break measured by ASTM D638-14. The terminal interference screw is deburred, inspected for knit lines at the thread root, and sterilized by ethylene oxide under ISO 11135:2014 with residual limits assessed according to ISO 10993-7:2008; gamma irradiation is avoided unless a validated low dose is shown not to reduce inherent viscosity below the device specification. In vitro degradation for absorbable internal fixation devices follows ASTM F1635-16 in phosphate-buffered saline at 37 °C and pH 7.4, where high-IV PLLA of this class typically retains the majority of its initial molecular weight during the first 6 months while mass loss occurs later.

Craniomaxillofacial Mesh Annealing After Thermoforming

Craniomaxillofacial fixation devices made from PURASORB PL 24 are produced by sheet extrusion followed by thermoforming for orbital floor and mandibular fracture applications. Sheet extrusion is performed on a single-screw extruder with a barrier screw and L/D of 24:1; die temperatures are maintained at 180 °C to 200 °C, and the melt is cast onto polished chill rolls at 15 °C to 25 °C to produce amorphous sheet with a thickness tolerance of ±0.02 mm across the forming area. Thermoforming is carried out at 90 °C to 120 °C, above the glass transition temperature of PLLA but below the melt crystallization onset; the formed mesh or plate is then annealed at 110 °C to 130 °C for 2 h to 6 h to raise crystalline content and reduce creep under screw compression. Differential scanning calorimetry according to ASTM D3418-15 is used to confirm that the annealed part reaches the target crystallinity range, which for oriented PLLA fixation devices is typically between 30% and 50%; higher crystallinity increases modulus but narrows the elongation at break to the lower end of the 2% to 6% range reported for PLLA homopolymers under ASTM D638-14. Compliance for CMF absorbable plates and screws is governed by ASTM F2502-17, while biological evaluation follows ISO 10993-1:2018 with implantation assessment under ISO 10993-6:2016; if the device is sterilized by ethylene oxide, ISO 11135:2014 and ISO 10993-7:2008 apply. The terminal CMF mesh or plate is trimmed, annealed, and inspected for microcracks at the strut intersections using low-magnification light microscopy; because PLLA has low impact toughness compared with titanium, bend angles during surgery are restricted by the device design validation. Blending PURASORB PL 24 with lower-viscosity PLLA or poly(L-lactide-co-glycolide) copolymer is used only when faster resorption is required, and the blend ratio is set by the target in vitro molecular weight retention curve generated under ASTM F1635-16.

What Limits Small-Shot Suture Anchor Moulding with a 2.4 dL/g PLLA Homopolymer?

For rotator cuff and glenoid labral anchors, PURASORB PL 24 is processed on micro-injection moulding machines with shot weights typically below 1 g; the high melt viscosity of a 2.4 dL/g PLLA homopolymer forces the barrel temperature to the upper end of the processing window, often 190 °C to 200 °C, which is close to the temperature where thermal degradation of the polyester backbone accelerates. Because the screw diameter in micro-moulders is usually 12 mm to 16 mm, surface-to-volume conditions are severe, and total melt residence time in the plastication unit and hot runner is held below 4 min; hot runner nozzles are set at 185 °C to 190 °C, and the gate is placed at the anchor body end rather than at the suture eye to avoid gate vestige acting as a stress concentrator. Formulation for suture anchors uses unplasticized PURASORB PL 24 or a blend with poly(L-lactide-co-glycolide) at 10 wt% to 25 wt% when a shorter resorption interval is required; the blend ratio is adjusted only after the copolymer ratio and L/G ratio are specified, because glycolide accelerates water uptake and changes the degradation profile. The terminal implant is conditioned at room temperature in dry nitrogen for at least 24 h before packaging to prevent moisture regain, and ethylene oxide sterilization is applied under ISO 11135:2014. The relevant material standard for in vitro degradation is ISO 13781:2017, with tensile and flexural properties on moulded specimens tested according to ASTM D638-14 and ASTM D790-17, while the implanted device is evaluated under ISO 10993-1:2018 and ISO 10993-6:2016 for local tissue response. The key process parameter is not injection speed alone but the ratio of shot volume to barrel capacity; if this ratio falls below 20%, the resin remains in the barrel for multiple thermal cycles, and molecular weight loss can be detected as a reduction in suture eye tensile strength after sterilization.

Guided bone regeneration membranes made from PURASORB PL 24 are produced by cast-film extrusion rather than injection moulding because the film must maintain thickness uniformity across a planar area and possess a smooth surface to limit epithelial attachment. The resin is vacuum-dried to below 250 ppm moisture and extruded through a flat die at 180 °C to 200 °C onto chill rolls at 15 °C to 20 °C; the resulting amorphous film is then optionally oriented in the machine direction at 70 °C to 80 °C to increase tear resistance without inducing excessive crystallinity that would make the membrane too stiff. For dental barrier indications, the device must maintain barrier function for 4 to 6 months and then undergo hydrolytic degradation; PURASORB PL 24 is selected because the high inherent viscosity extends the period before significant mass loss occurs, whereas a lower-viscosity PLLA or a PLGA copolymer would lose mechanical integrity earlier. Residual lactide monomer in the final film is controlled by vacuum extraction before packaging, and the monomer content is measured by an internal HPLC method; elevated residual lactide contributes to local pH depression during resorption and is not controlled by the film extrusion step alone. Biological evaluation follows ISO 10993-1:2018, with cytotoxicity under ISO 10993-5:2009 and sensitization or irritation under ISO 10993-10:2010; the film is sterilized by ethylene oxide according to ISO 11135:2014, and residual EO is verified against ISO 10993-7:2008. The terminal membrane is die-cut into rectangular or trapezoidal shapes, inspected for pinholes using transmitted light, and sealed in foil laminate with desiccant because moisture uptake during storage can initiate pre-implantation hydrolysis.

When a Permanent Polymer Is Replaced by PLLA in Bioresorbable Vascular Scaffolds

In tubular form, PURASORB PL 24 is evaluated for fully bioresorbable vascular scaffolds, where the polymer must survive extrusion, tube drawing, laser cutting, expansion, and crimping without losing the molecular weight needed for radial support during the first 3 to 6 months after implantation. The tube is produced on a single-screw extruder with melt temperatures of 190 °C to 200 °C, drawn at 70 °C to 90 °C to induce orientation, and then laser-cut into a strut pattern; the orientation step raises hoop strength, but the scaffold design must account for the low elongation at break of PLLA, which is typically 2% to 6% under ISO 527-2:2012. Because PLLA is radiolucent, scaffold visualization during implantation is limited unless a radiopaque marker is added at the ends. The critical processing boundary is the reduction in molecular weight that occurs during thermal processing; if the inherent viscosity of the cut scaffold drops below a device-specific threshold, the strut radial force and fatigue life measured under cyclic loading decrease. Published data for this specific configuration of PURASORB PL 24 in vascular scaffolds is limited, so development programs use ASTM F1635-16 in vitro degradation testing and ISO 25539-2:2020 endovascular device evaluation to generate device-specific material qualification rather than relying on orthopaedic degradation data. Biological evaluation follows ISO 10993-1:2018 with systemic toxicity under ISO 10993-11:2017 and implantation under ISO 10993-6:2016; sterilization is by ethylene oxide or, in early feasibility evaluations, by electron beam only at doses validated to keep molecular weight loss within specification. The terminal scaffold struts are inspected under scanning electron microscopy for laser-induced edge defects, and any microcrack at the strut intersection is cause for rejection because stress concentration during crimping and balloon expansion can trigger brittle fracture.

Drug-Eluting Depot Matrices and PLLA Inherent Viscosity Effects on Burst Release

Because high inherent viscosity lowers matrix diffusivity and delays the onset of the erosion phase, PURASORB PL 24 is selected for injectable microspheres and pre-formed drug-eluting depots. A solvent evaporation process is used for microsphere manufacture: the polymer is dissolved in dichloromethane at a polymer concentration of 3% to 10% w/v, the active pharmaceutical ingredient is dispersed or co-dissolved at a drug-to-polymer ratio between 1:4 and 1:20 depending on solubility, and the organic phase is emulsified into an aqueous polyvinyl alcohol continuous phase; the solvent is then removed under controlled stirring to produce microspheres with a mean particle size controlled by impeller geometry and continuous-phase viscosity. The terminal depot formulation is washed, vacuum-dried to below 0.5% residual dichloromethane, and packaged under nitrogen; residual solvent limits are set by ICH Q3C guidelines, not by the polymer standard. Compliance as a drug-device combination product is more complex than for a polymer-only implant: the polymer is qualified under ISO 13485:2016 and ISO 10993-1:2018, but the finished microsphere or implant is tested under the pharmacopoeial particle counting method USP <788>, leachables protocols, and the applicable degradation method ASTM F1635-16 for the hydrolytic matrix. The high molecular weight of PURASORB PL 24 delays the onset of the erosion phase, so early release is governed mainly by diffusion through the amorphous regions; if the polymer is annealed before or during processing, increased crystallinity reduces the initial burst but can also reduce total drug diffusivity by an amount that must be determined experimentally. The main processing incompatibility is overheating of the solvent-containing solution during solvent removal; if the solution temperature is raised above 40 °C, the polymer may precipitate asymmetrically and create a skin layer that changes the release profile.