PURASORB PDLG 75016A is supplied as a dry, acid-terminated poly(D,L-lactide-co-glycolide) with a nominal 75:25 molar ratio and an inherent viscosity midpoint of 0.16 dL/g measured in chloroform at 25°C at 0.1 g/dL. The acid-terminated chain ends are quantified on the certificate of analysis by titration and are selected for long-acting injectable microspheres because they increase water uptake and accelerate bulk hydrolysis. A double-emulsion solvent-evaporation process is used for peptide-loaded depot formulations. The organic phase is prepared as 180–250 mg/mL polymer in dichloromethane. The inner aqueous phase contains the peptide at 5–10% w/v and may include trehalose or mannitol at 5–10% w/v as a lyoprotectant. Primary emulsification is performed with a Silverson L5M-A rotor-stator equipped with a square-hole high-shear screen at 6,000–10,000 rpm for 60–120 s. The primary water-in-oil emulsion droplet size is typically 1–5 µm by optical microscopy before transfer. The resulting emulsion is transferred into a continuous aqueous phase containing 0.5–2.0% w/v poly(vinyl alcohol) at 2–8°C, followed by secondary emulsification at 2,500–4,000 rpm. Solvent extraction and matrix hardening proceed in an aqueous bath at 15–25°C with a water-to-organic phase ratio of at least 10:1 v/v. Hardened microspheres are retained on a 20 μm stainless-steel sieve and collected below an 80 μm sieve, washed with water for injection, and lyophilized. A prolonged hold above 20°C beyond 8 h during hardening has been observed on production batches to cause surface pitting and a reduction in weight-average molecular mass of more than 10%, due to autocatalytic hydrolysis of the free carboxyl chain ends. Residual dichloromethane after vacuum drying at 25°C for 24 h is typically below 100 ppm, and the release limit is 600 ppm per ICH Q3C. Residual poly(vinyl alcohol) determined by size-exclusion chromatography is limited to 2.0% w/w. In-vitro release testing uses USP Apparatus 4 flow-through cells with phosphate-buffered saline pH 7.4 at 37°C. Representative batches of this acid-terminated grade show an initial 24 h release of ≤10% and a lag phase of 2–5 days before erosion-controlled release; this lag phase is shorter than the corresponding ester-terminated 75:25 grade under identical process conditions. The final microsphere suspension is assessed according to ISO 10993-1:2018 subacute implantation, USP <71> sterility, USP <85> bacterial endotoxins, and USP <788> particulate matter. Sterilization is by gamma irradiation at 25–40 kGy because steam autoclaving induces bulk polymer degradation before sterility can be achieved. Particle size is measured by laser diffraction per ISO 13320:2020, with a target volume mean diameter of 30–70 μm for intramuscular injection.

When Acid-Terminated PLGA Is Dissolved in N-Methyl-2-Pyrrolidone for In Situ Depot Formation

For subcutaneous in situ-forming depot systems, PURASORB PDLG 75016A is dissolved in N-methyl-2-pyrrolidone at 35–50% w/w under low-shear agitation for 12–24 h at 25°C. The drug substance is then incorporated by planetary mixing at 25 rpm for 16 h in a 2 L Teflon-coated anchor mixer. The resulting solution viscosity at 25°C is 3,000–6,000 mPa·s at 10 s⁻¹ as measured by cone-and-plate rheometry. The solution is filled into a 5 mL glass syringe with a 21-gauge needle. Upon injection into subcutaneous tissue, NMP exchanges with aqueous tissue fluid and the polymer precipitates to form a coherent depot. The higher carboxyl end-group density of this acid-terminated grade increases the initial water uptake of the precipitated matrix, and comparative release data show a shortening of the induction phase by 1–3 days relative to ester-terminated 75:25 PLGA at equivalent polymer concentration. Polymer concentrations below 30% w/w produce fragmented or particulate depots rather than a coherent implant, and burst release exceeds 30% of drug in the first 12 h. At NMP concentrations above 60% w/w, rapid solvent efflux collapses the outer skin and similarly increases burst release. The formulation is therefore maintained within 35–50% w/w polymer. Filtration through a 0.2 μm sterilizing-grade PTFE membrane is not feasible because of the high solution viscosity. Terminal gamma irradiation at 25 kGy is used, but irradiation of NMP solutions generates free radicals that reduce polymer molecular mass by an additional 5–8%; this reduction is accounted for in release design. Residual NMP in the implanted matrix falls below 1% w/w after 7 days in vivo based on published phase-inversion depot studies. The prefilled syringe system is assessed under USP <790> visible particulates, USP <788> subvisible particulate matter, ISO 11040-4 prefilled syringes, and ISO 10993-6:2016 implantation testing. The final product is a sterile, single-use prefilled subcutaneous depot.

A solvent-displacement route for nanoparticle formation starts with PURASORB PDLG 75016A dissolved at 5–10 mg/mL in acetonitrile or acetone. The organic solution is introduced dropwise into an aqueous stabilizer phase containing 0.1–0.5% w/v poloxamer 188 or D-α-tocopheryl polyethylene glycol 1000 succinate under magnetic stirring at 400–800 rpm. Rapid diffusion of the water-miscible solvent produces nanoparticles with a z-average diameter of 90–180 nm and a polydispersity index of 0.08–0.15 measured by dynamic light scattering at 25°C according to ISO 22412:2017. At pH 7.4, the carboxylic acid end groups are ionized, giving the particle surface a zeta potential of −25 to −45 mV in 1 mM KCl. For ligand conjugation, the terminal carboxyl groups are reacted with 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide and N-hydroxysuccinimide at a molar ratio of 5:1 to carboxyl sites in 0.1 M MES buffer at pH 5.5. Residual acetonitrile is removed by rotary evaporation at 30°C under 200 mbar and by nitrogen stripping; the drug product limit is 410 ppm under ICH Q3C. Unencapsulated drug is removed by tangential flow filtration on a 100 kDa regenerated cellulose membrane, yielding entrapment efficiency above 70% for hydrophobic actives. Lyophilization is performed with sucrose or trehalose at 5–10% w/w using a cycle with primary drying at −30°C shelf temperature and 50 mTorr chamber pressure for 24 h, followed by secondary drying at 20°C for 6 h. In-vitro release in phosphate-buffered saline pH 7.4 at 37°C shows a biphasic profile: 5–15% release in the first 24 h followed by sustained release over 10–20 days. The finished nanoparticle product is intended for intravenous or intratumoral administration and is tested according to ISO 10993-5:2009 cytotoxicity, USP <85> endotoxins, and USP <788> particulate matter. The process is limited to 2 L batch volumes in a flameproof enclosure because solvent flammability and adiabatic heating in high-shear homogenizers accelerate polyester degradation.

How Does Hot-Melt Extrusion Avoid Solvent Residuals in Subdermal Rod Manufacture?

On a vented twin-screw extruder with a 16 mm screw diameter and an L/D ratio of 25:1, PURASORB PDLG 75016A is melt-processed into subdermal rods. Pre-drying is performed at 30°C and 5 mbar for 24 h because acid-terminated PLGA undergoes autocatalytic hydrolysis if moisture exceeds 500 ppm. The feed zone is set at 60°C, the mixing zone at 75°C, and the die at 80°C; melt temperature at the die is 78–85°C. Screw speed is limited to 50–100 rpm to prevent shear heating, and motor torque remains below 4 N·m. Residence time is kept below 120 s. At die temperatures above 100°C, the weight-average molecular mass loss exceeds 10% in a single pass and the extrudate becomes discolored. The drug substance is pre-blended with a portion of the polymer and fed as a physical mixture; loadings of 10–25% w/w process without phase separation. At loadings above 30% w/w, the melt viscosity falls below 200 Pa·s and strand dimensional stability is lost. A circular die of 1.2 mm diameter produces rods that are cooled on a 4°C stainless-steel belt and cut to 5–10 mm lengths. The rods are annealed at 40°C for 12 h under dry nitrogen to reduce residual stress. The final subdermal rods are tested under ISO 13781:2017 for lactide-based surgical implant polymers, USP <905> uniformity of dosage units, and USP <788> particulate matter. In-vitro release in phosphate-buffered saline pH 7.4 at 37°C from a 1.2 mm rod with 20% w/w drug loading shows an initial 24 h release below 10% and near-zero-order release over 14–21 days. The principal production failure mode is strand breakage when feed continuity is interrupted; batch-to-batch inherent viscosity variation of ±0.02 dL/g can shift die pressure by 10–15% and requires feed-rate compensation of 5–10%.

Solvent-cast bioresorbable films for periodontal pocket insertion are prepared by dissolving PURASORB PDLG 75016A at 10–15% w/w in acetone or ethyl acetate under rolling agitation at 25°C for 6–12 h. The solution is cast onto a silicone-coated polyester release liner using a knife-over-roll coater with a wet gap of 0.5–1.0 mm. Drying is performed in a two-zone forced-air oven: the first zone at 25°C for 30 min to limit skin formation, and the second zone at 35°C for 2–4 h under reduced pressure to drive off residual solvent. The dried film thickness is 150–300 μm. The acid-terminated grade yields films with higher water permeability than ester-terminated PLGA of equivalent inherent viscosity; this property is used to control drug release in gingival crevicular fluid. Uniformity of drug content is evaluated by USP <905>. Mechanical properties are determined by tensile testing per ASTM D882-18; elongation at break typically remains above 5% at 23°C and 50% RH. The residual ethyl acetate limit is 5000 ppm under ICH Q3C. Films are cut to 10 mm × 10 mm or 5 mm × 20 mm and packaged in nitrogen-purged aluminum foil. Placement into the periodontal pocket is intended for local delivery of an antimicrobial agent over 7–14 days. In-vitro dissolution testing uses USP Apparatus 5 paddle-over-disk in synthetic gingival crevicular fluid at 37°C. The final product is assessed under ISO 10993-5:2009 cytotoxicity and ISO 10993-10 irritation testing; sterility may be by gamma irradiation at 25 kGy if the drug substance is radiation-stable.

Drug-Eluting PLGA Coatings for Orthopaedic Trauma Hardware: Adhesion and Release

For orthopaedic trauma hardware, PURASORB PDLG 75016A is dissolved at 1–2% w/v in ethyl acetate or acetone to form an ultrasonic spray-coating solution. Titanium or stainless-steel screws are cleaned with 70% isopropanol and dried with nitrogen before coating. An ultrasonic nozzle operating at 60 kHz with a flow rate of 0.5–1.0 mL/min deposits alternating polymer and drug-containing layers onto the implant surface held at 40–60°C. The target coating thickness is 5–15 μm. Drug loading is 5–10% w/w relative to polymer mass. The coated screws are annealed at 40°C for 24 h to remove residual solvent and to improve adhesion. Coating adhesion is assessed by a tape test at room temperature per ASTM D3359-23; for implant use, additional scoring and peel testing are performed according to the sponsor's coating qualification protocol. Residual ethyl acetate is limited to 5000 ppm under ICH Q3C. Release testing in phosphate-buffered saline pH 7.4 at 37°C from a 10 µm coating with 10% w/w drug loading typically shows release over 21–35 days, with an initial 24 h release of 10–20%. The coated implant is assessed for biocompatibility under ISO 10993-6:2016 implantation and ISO 10993-5:2009 cytotoxicity. The primary failure mode is delamination at the metal-polymer interface when the implant is subjected to insertion torque above 4 N·m; published data for this specific low-IV acid-terminated grade on orthopaedic hardware is limited, so adhesion under clinical insertion must be verified on the final geometry.