PURASORB PL 38 is taken into load-sharing osteosynthesis programmes as a poly(L-lactide) homopolymer with a nominal inherent viscosity of 3.8 dL/g specified against chloroform at 25 °C. In cortical screw and plate lines, the formulation is intentionally kept at 100 wt% as-supplied resin because every additional low-molecular-weight species—plasticisers, lubricants, or unbound nucleants—changes the hydrolytic degradation rate and the retention of bending strength after implantation. Residual L-lactide monomer is controlled below 0.5 wt% on the batch certificate using methods aligned with ISO 13781:2017, and the same certificate reports tin catalyst residues because these residues participate in the degradation kinetics of the finished device. The injection molding feed therefore is 100 parts PLLA with no filler; if a coloured or laser-markable variant is required, the masterbatch is limited to 0.05–0.10 wt% and must be covered by the device master file. Manufacturing is conducted under ISO 13485:2016, and the biological evaluation of the finished osteosynthesis device follows ISO 10993-1:2018. Before melt processing, the resin is vacuum-dried at 80 °C until the residual moisture is below 250 ppm by Karl Fischer. Production lines that omit this step observe splay and bubble defects in machined blanks, because steam volatilisation at the screw tip creates voids that act as stress concentrators after implantation. The material is fed into an all-electric injection molding machine with closed-loop switchover from velocity to pressure control. Melt temperatures are maintained between 190 °C and 215 °C; below 185 °C unmelted gel particles survive, and above 220 °C thermal scission reduces molecular weight within a single barrel residence time. For 2.5 mm-thick fracture plates, cavity pressures observed on multi-cavity tools are in the range 1000–1400 bar. After ejection, the amorphous parts are annealed in nitrogen at 110–120 °C for 4–8 h to increase crystallinity from less than 10% to 35–45% as measured by differential scanning calorimetry. This step is critical: insufficient annealing leaves screws with creep under shear loading; excessive annealing embrittles sharp thread crests and produces tip fracture during insertion torque testing. The terminal products are bioresorbable screws, pins, and small plates for non-load-bearing or load-sharing fracture regions such as the ankle syndesmosis, elbow, and hand. The devices are sterilised by ethylene oxide because steam autoclave conditions at 121 °C for 15 min initiate bulk hydrolysis and warp thin plates beyond dimensional tolerance. Gamma irradiation above 25 kGy is generally avoided because it can reduce molecular weight and shift the mechanical retention profile; where radiation is selected, this exposure must be included in the design verification matrix.

What Limits the Injection Molding Window for High-IV PLLA Suture Anchors?

Arthroscopic shoulder and knee reconstruction lines convert PURASORB PL 38 into interference screws and suture anchors with a deliberate balance between high inherent viscosity and flow length. The formulation for ligament fixation screws is often 100 wt% high-IV PLLA; for anchors placed in cancellous bone, 15–30 wt% β-tricalcium phosphate is compounded into the polymer to buffer acidic degradation products and create an osteoconductive interface. Above 30 wt% filler, the notched impact strength drops rapidly and the anchor eyelets crack during deployment. The filler is surface-treated before compounding to prevent moisture uptake at the filler-polymer interface. The unfilled resin is evaluated under ASTM F1925-17, while the composite device falls under ASTM F2502-17 for absorbable polymeric implants. Multi-cavity tools for 3.0 mm to 5.0 mm interference screws are run at melt temperatures of 190–210 °C, with water-cooled mold temperatures of 25–40 °C to maintain an amorphous, low-shrinkage part. The process conflict is that high-IV PLLA has a low melt flow index; therefore, gates must be placed on the drive end and wall thickness kept below 3.5 mm to avoid sink marks. Screw recovery time on a 40 mm plasticizing screw can exceed 30 s per shot, and this extended residence time is a known source of batch-to-batch inherent viscosity drift. Production lines mitigate this by reducing screw back pressure to 5–10 bar and using a low-compression screw profile. Post-mould annealing at 110 °C for 4 h under vacuum raises crystallinity and stabilises thread geometry. The finished devices are threaded interference screws for ACL reconstruction, push-in suture anchors for rotator cuff repair, and ribbed fixation nails for ligament repair.

Craniomaxillofacial Plate Thermoforming and the Residual-Monomer Constraint

Craniomaxillofacial surgeons use absorbable PLLA plates and screws in paediatric vault reconstruction and orbital floor repair, where a second removal surgery is undesirable. The plates are produced from compression-moulded PLLA sheet at 100 wt% resin; no solvent, plasticiser, or additional monomer is introduced. Residual lactide monomer is a release concern in the oral cavity, so the sheet is controlled to less than 0.5 wt% according to ISO 13781:2017, and the cytotoxicity endpoint of ISO 10993-5:2009 is repeated on the sterilised device. Sensitisation testing is performed under ISO 10993-10:2010. The terminal product types include straight and L-shaped microplates, screw diameters from 1.5 mm to 2.5 mm, and orbital floor sheets with thicknesses from 0.3 mm to 1.0 mm. Sheet is compression-moulded at 190–200 °C under 50–100 bar, then water-cooled at 10–15 °C/min to maintain a low-crystallinity state that permits intraoperative contouring. The surgeon heats the plate in a sterile water bath at 55–70 °C and bends it to the bone contour; below 55 °C the PLLA sheet remains too stiff for manual shaping, and above 70 °C the plate can lose edge definition. CNC machining after molding produces the screw holes and plate outline; ultrasonic cleaning in isopropanol and final ethylene oxide sterilisation complete the line. The process boundary is that dry machining creates sheared edges with microcracks; therefore, cutting tools are water-lubricated and feed rates are limited to avoid local temperatures above the glass transition.

Bioresorbable vascular scaffolds represent a different conversion route: the PLLA is not injection-moulded into a discrete part but is extruded into a tube that must maintain dimensional accuracy over a length of 80–150 mm. In this application, PURASORB PL 38 forms the backbone at 100 wt% because semi-crystalline PLLA provides higher radial strength than amorphous PDLLA. The anti-proliferative coating is a separate layer of PDLLA with a drug-to-polymer ratio fixed in the pharmaceutical development report; published scaffold evaluations report drug loadings of 10–30 wt% relative to coating mass. The scaffold is evaluated under ISO 25539-2:2020 for vascular stents and ASTM F3067-23 for absorbable stent characterisation, with hemocompatibility per ISO 10993-4:2017. Because the tube is the critical durability component, the inner and outer diameter tolerances are held to ±0.05 mm, and wall thickness is controlled to 0.10–0.20 mm. Tube extrusion is run at 190–220 °C with nitrogen purge on the feed hopper; the melt is drawn down over a precision mandrel and quenched in a water bath at 15–20 °C to lock in an amorphous structure. A second expansion step at 80–90 °C introduces strain-induced crystallinity and improves radial strength. Laser cutting with a 355 nm UV source then creates the scaffold geometry; post-cutting rinsing removes debris before crimping onto a delivery balloon. The principal production failure mode is ovality caused by non-uniform draw-down; ovality above 1% leads to crimp asymmetry and unacceptable deployment geometry. Ethylene oxide sterilisation is preferred over gamma because high-energy radiation can reduce PLLA molecular weight and shorten the load-bearing period below the intended 6–12 months. The terminal product is a bioresorbable coronary scaffold, not a permanent stent, and is contraindicated for heavily calcified lesions where radial force is insufficient.

Soft-tissue volumizing injection products use PLLA in a different physical form: sterile, resorbable microspheres rather than load-bearing solid parts. A representative injectable presentation is 150 mg PLLA microspheres per vial, lyophilized with mannitol and sodium carboxymethylcellulose as processing excipients; after reconstitution with 5–10 mL sterile water for injection, the suspension is injected into the deep dermis or subcutis. The PLLA fraction is therefore 100 wt% of the particle, with the excipients removed or remaining at low levels in the final device. Particle size is controlled to 40–63 μm by sieving because particles below 20 μm are more likely to be phagocytosed and may trigger granulomatous responses. Microspheres are produced by oil-in-water solvent evaporation. A 5–10 wt% PLLA solution in dichloromethane is emulsified into an aqueous phase containing 1–5 wt% poly(vinyl alcohol); the solvent is removed under vacuum at 35–40 °C. The hardened spheres are washed to reduce PVA residues below 0.5 wt%, lyophilized under aseptic conditions, and filled in an ISO 14644-1:2015 class 5 environment. The process is governed by ISO 10993-1:2018 for biological evaluation and ISO 10993-6:2016 for implantation effects. Terminal product types are injectable poly-L-lactic acid implants for facial volume restoration, including cheek and temple hollows; superficial intradermal placement is associated with nodule formation and is avoided.

When Semicrystalline PLLA Enters an Electrospinning Train

Tissue-engineering programmes that need fibre-based PLLA scaffolds dissolve the polymer into a volatile solvent system rather than using melt processing. A common electrospinning solution is 8–12 wt% PLLA in chloroform: N,N-dimethylformamide 80:20 vol/vol; the solution viscosity must be high enough to prevent Rayleigh breakup and low enough to pass through a 0.45 μm filter before the spinneret. The solution is metered at 0.5–2.0 mL/h through a 22-gauge blunt-tip needle at a tip-to-collector distance of 12–18 cm and a potential difference of 15–25 kV. A rotating mandrel at 500–2000 rpm produces aligned fibres with diameters from 300 nm to 1.5 μm. After deposition, the scaffold is dried under vacuum at 40–50 °C to reduce residual chloroform and DMF below the leachables limits of ISO 10993-18:2020. Cytotoxicity is evaluated under ISO 10993-5:2009. The terminal products are electrospun PLLA membranes and tubular scaffolds for tendon repair, vascular tissue engineering, and guided bone regeneration. The key limitation is mechanical strength: an electrospun mat has an effective tensile modulus orders of magnitude lower than injection-moulded PLLA, so it is not a load-bearing structure. Below 8 wt%, bead defects form; above 12 wt%, the spinneret blocks due to solvent evaporation at the tip.