Orthopaedic Interference Screws and Suture Anchors

Melt compounding of RESOMER L 207 S for orthopaedic fixation begins with vacuum drying at 80 °C for a minimum of 4 h to reduce moisture below 250 ppm, because hydrolytic chain scission during plastication lowers inherent viscosity and shifts the molecular weight distribution toward brittle failure. On production-scale reciprocating screw injection moulding machines, barrel profiles from feed to nozzle are maintained between 170 °C and 205 °C, with screw L/D ratios of 18:1 to 24:1 and low compression ratios of 2.0:1 to 2.5:1. A cold mould set at 20–40 °C freezes the amorphous matrix rapidly and prevents premature spherulitic growth that produces warpage in interference screws with thread depths above 0.8 mm. Post-mould annealing at 110–125 °C for 2–6 h raises the degree of crystallinity into the 40–55 % range, which slows in vivo strength loss and improves dimensional stability at the screw-bone interface. Injection pressure is typically held between 80 MPa and 140 MPa, and hold pressure is maintained until gate freeze to avoid sink marks in thick-walled screw heads. The terminal devices include knee and shoulder interference screws, suture anchors with distal eyelets, and craniofacial tacks. Resin compliance is controlled under ASTM F1925-17 for virgin semi-crystalline poly(L-lactide), and finished implant mechanical testing follows ASTM F2502-17 for absorbable plates and screws. Biological evaluation is conducted under ISO 10993-1:2018 with implantation testing under ISO 10993-6:2016; in vitro degradation is profiled according to ISO 13781:2017 in phosphate-buffered saline at 37 °C. A production bottleneck occurs when barrel residence time exceeds 15 min at melt temperatures above 210 °C, because intramolecular transesterification generates lactide monomer and depresses melt viscosity. Injection gate diameters below 0.5 mm for high-viscosity PLLA can produce shear heating above 10,000 s⁻¹ and localised yellowing. The practical processing window for this grade is therefore bounded by moisture content, barrel temperature, and gate design rather than by screw speed alone.

Solvent-based encapsulation converts the high-molar-mass PLLA into drug-loaded microspheres in which the polymer solution is dispersed in an aqueous continuous phase and the solvent is stripped under controlled extraction. In a typical oil-in-water process, RESOMER L 207 S is dissolved in dichloromethane at 5–15 % w/v, and the active pharmaceutical ingredient is either co-dissolved or suspended at drug-to-polymer ratios from 1:5 to 1:20. The organic phase is emulsified into 0.1–1.0 % w/v polyvinyl alcohol solution under rotor-stator homogenisation at 5,000–20,000 rpm, producing a pre-hardened droplet population with a median size between 10 µm and 100 µm. Solvent extraction is performed by transferring the emulsion to a stirred aqueous bath at 15–25 °C for 3–6 h, so that dichloromethane partitions into the continuous phase without forming a collapsed surface skin. Hardened microspheres are collected by filtration, washed, and vacuum-dried at 30–40 °C to residual solvent levels below 600 ppm as required for dichloromethane under ICH Q3C(R8). Terminal products in this segment include injectable depot suspensions for peptide delivery and implantable subcutaneous rods for long-term hormone suppression. Release kinetics are governed by PLLA hydrolysis, with an initial diffusion-controlled phase followed by autocatalytic ester cleavage inside the polymer matrix; published data for this specific configuration is limited, but reproducible release windows typically span 30–180 days depending on particle size and molecular weight. The relevant quality standards include USP <711> dissolution testing, ISO 10993-1:2018 biological evaluation, and aseptic manufacture under 21 CFR 210 and 21 CFR 211. A critical process conflict arises when the aqueous continuous phase temperature is raised above 25 °C during solvent extraction, because dichloromethane evaporation within the droplet creates internal voids and lowers core encapsulation efficiency. Likewise, homogenisation above 20,000 rpm generates viscous heating and broadens the particle size distribution, while insufficient continuous-phase surfactant below 0.05 % w/v permits droplet coalescence and agglomeration. The process is therefore operated with tight control of phase ratio, shear rate, and extraction temperature rather than by fixed agitation time alone.

Why Does Molecular Orientation Govern Monofilament Tensile Properties?

Melt-spun PLLA monofilament exits the spinneret as a mostly amorphous strand with low tensile strength, and the clinical utility of the fibre is created only after controlled orientation and heat setting. RESOMER L 207 S is dried to below 200 ppm moisture and extruded through a single-screw extruder with a die land length of 20–30 mm at melt temperatures from 185 °C to 205 °C. The as-extruded filament is quenched in air at 10–25 °C and then passed through heated godets at 60–90 °C to impose draw ratios between 5:1 and 10:1. Drawing aligns amorphous chain segments along the fibre axis and induces strain-induced crystallisation, producing a fibrillar microstructure with higher modulus and reduced elongation at break. Published values for oriented PLLA monofilaments typically report tensile strengths above 400 MPa, moduli above 4 GPa, and knot-pull strength retention sufficient for resorbable suture and ligament fixation applications. The terminal products are monofilament sutures, suture anchors, and oriented fibre-reinforced plates or rods in which the oriented phase carries load while the surrounding PLLA matrix degrades slowly. Conformance testing uses ISO 13781:2017 for degradation profiling, ASTM D2256-21 for tensile properties of yarns, and ISO 10993-1:2018 for biocompatibility. A specific failure mode in production is filament breakage when the draw ratio exceeds 10:1 before the godet temperature reaches 60 °C, because the oriented amorphous fraction cannot relax and fibrillation occurs. A second constraint is that residual moisture above 250 ppm in the dried pellets produces microvoids in the extrudate that lower drawability and create surface roughness on the monofilament. The orientation window is therefore bounded at the lower end by insufficient chain mobility and at the upper end by strain hardening and filament rupture.

In balloon-expandable scaffold manufacturing, the extruded PLLA tube is radially expanded and laser-cut into a strut network before crimping onto a delivery balloon. RESOMER L 207 S is first melt-extruded into tubing with an outer diameter of 1.5–2.0 mm at 180–200 °C, then conditioned at 70–90 °C under positive internal pressure to induce biaxial orientation. Radial expansion produces hoop-direction orientation that resists compressive loading in the implanted device, while axial draw during tube processing controls foreshortening. The oriented tube is fixed by rapid cooling to below the glass transition, then cut with a femtosecond laser system at pulse durations below 500 fs to produce strut widths of 100–200 µm without significant heat-affected zones. The scaffold is crimped at 37 °C onto a delivery catheter using multi-segment radial crimpers, then sterilised by ethylene oxide rather than gamma irradiation because ionising radiation causes chain scission and molecular weight loss in PLLA. Terminal products are bioresorbable coronary scaffolds designed to provide temporary vessel support during the 6–12 month revascularisation window. Published data for this specific configuration is limited, and manufacturers rely on radial strength and recoil acceptance criteria defined in design verification plans under ISO 25539-2:2020 rather than a single published threshold. Regulatory evaluation references ISO 25539-2:2020 for vascular stents and ISO 10993-1:2018 for biocompatibility, with degradation testing conducted under ISO 13781:2017 at 37 °C in phosphate buffer. The dominant production bottleneck is strut fracture during crimping, because oriented PLLA has high strength in the hoop direction but limited plastic deformation capacity. Laser-induced microcracks propagate when the crimping angle exceeds 30° at room temperature, and the defect rate rises when the scaffold is stored above 25 °C before deployment. Therefore the crimping step is conducted in a temperature-controlled environment with humidity held below 30 % RH, and the strut pattern is designed to avoid acute angles below 90° at the node junctions.

When PLLA Filament Is Deposition-Printed for Patient-Specific Craniofacial Scaffolds

Fused filament fabrication of RESOMER L 207 S requires filament dried to below 250 ppm moisture, nozzle temperatures of 200–220 °C, and a build plate at 55–70 °C. Layer heights of 0.10–0.25 mm and print speeds of 20–60 mm/s produce patient-matched cranial defect templates and maxillofacial bone void fillers, but interlayer weld strength remains 40–60 % lower than in-plane strength when the nozzle temperature falls below 195 °C. The terminal constructs are evaluated under ISO/ASTM 52900:2021, ASTM D638-14, ASTM D695-15, and ISO 10993-5:2009. Retained moisture above 300 ppm creates porosity and reduces molecular weight during extrusion.

Electrospinning of the resin into submicron fibres produces high-porosity membranes used as guided tissue regeneration barriers and adhesion prevention sheets. RESOMER L 207 S is dissolved in a mixed solvent system consisting of chloroform and dimethylformamide at volume ratios of 70:30 to 85:15, with total polymer concentration of 6–10 % w/v. The solution is metered from a stainless-steel capillary at 0.5–2.0 mL/h into an electric field of 15–25 kV, while the collector distance is held at 10–20 cm. Fibre diameters from 300 nm to 1,200 nm are obtained by adjusting solution conductivity, flow rate, and relative humidity below 35 % RH. The resulting nonwoven membrane has pore diameters below 5 µm, which blocks gingival epithelial cell migration but permits diffusion of nutrients and metabolic waste. Terminal products include resorbable dental barrier sheets, periodontal regeneration membranes, and anti-adhesion films for tendon and peritoneal applications. The membranes are typically annealed at 80–100 °C for 1–2 h to increase crystallinity and reduce tear propagation under suturing. Compliance testing for this segment includes ISO 10993-1:2018 biological evaluation, ASTM F2902-16 for absorbable implants, and ISO 13781:2017 for in vitro degradation. A process conflict occurs when ambient relative humidity exceeds 40 % RH, because water vapour in the spinning chamber promotes rapid solvent evaporation and produces beaded fibres with weak mechanical interlocking. Conversely, a polymer concentration below 5 % w/v produces electrosprayed droplets rather than continuous fibres, while concentration above 12 % w/v raises viscosity to levels that prevent stable jet formation through a 21-gauge needle. The electrospinning environment is therefore controlled for humidity, solvent ratio, and solution conductivity to maintain a stable Taylor cone and uniform fibre populations.