In solvent evaporation–based microencapsulation of aqueous peptide payloads, 65:35 DL-PLG (B6001-2) is dissolved into dichloromethane at 50–300 mg/mL prior to formation of a water-in-oil primary emulsion. The primary emulsion is generated with a rotor–stator homogenizer at 10,000–15,000 rpm for 60–180 s before transfer into an outer aqueous continuous phase containing 0.5–2.0% w/v poly(vinyl alcohol) and 0.9% w/v sodium chloride. Solvent extraction is conducted in 2–5% aqueous isopropanol at 15–25 °C for 4–6 h, with residual dichloromethane reduced below the ICH Q3C class 2 limit of 600 ppm as measured by USP <467> headspace gas chromatography. Microspheres are collected on a 45 µm sieve, washed with water for injection, and vacuum-dried below <40 °C until residual moisture by Karl Fischer titrimetry is <1.0% w/w. Drug loading for leuprolide acetate and octreotide acetate depot systems commonly falls between 3–12% w/w, with encapsulation efficiency above 80% achievable only when the inner aqueous phase is buffered with acetate or phosphate and the osmotic pressure differential is controlled. Terminal sterilization with gamma irradiation at 25 kGy per ISO 11137-1 can reduce inherent viscosity by 20–50% depending on dose rate; aseptic filtration of the polymer phase through 0.22 µm PVDF membranes is therefore preferred for thermolabile peptides. Release testing follows USP <711> apparatus 2 or 4 in phosphate-buffered saline at 37±0.5 °C, where the 65:35 amorphous matrix typically hydrates within 24–48 h and exhibits bulk-erosion-controlled release over 4–8 weeks for low-solubility peptide salts. Subvisible particulate counts are controlled to USP <787> limits for injection volumes above 100 mL equivalent. The acid-terminated grade accelerates ester hydrolysis relative to ester-capped PLGA at equivalent lactide-glycolide ratio, which must be accounted for when setting release specifications for acid-labile actives.

What Limits Burst Release in Subcutaneous Depot Formulations?

Burst release in 65:35 DL-PLG depots is governed by pore coalescence, surface-associated drug, and early water ingress rather than by crystalline tortuosity, because the 65:35 copolymer is amorphous with a glass transition typically near 40–50 °C under dry conditions. When peptide loading exceeds 10% w/w, interconnected surface porosity formed during solvent extraction allows capillary wetting and dissolution of drug within the first 24 h; release may then exceed 30% of load before the first sampling interval. Reducing primary emulsion droplet size below 1 µm and lowering the solvent removal temperature to <20 °C densifies the surface skin and lowers burst release. Outer aqueous PVA above 2.0% w/v produces a viscous film that slows solvent flux but leaves residual poly(vinyl alcohol), which interacts with cationic peptides and alters encapsulation. Acid-terminal 65:35 DL-PLG generates an acidic microclimate during hydrolysis; for acid-labile payloads, addition of 3–10 wt% poorly soluble antacid excipients such as magnesium hydroxide moderates pH but also increases osmotic pressure inside the microsphere. Burst release is measured after 1 h and 24 h in phosphate-buffered saline at pH 7.4, with acceptance criteria often set at <20% cumulative release at 24 h for sustained-release injectable products. The exact burst value must be determined empirically for each peptide-polymer pair because it depends on drug solubility, primary emulsion stability, and residual solvent level at the microsphere surface.

For nanoparticle constructs intended for parenteral oncology applications, 65:35 DL-PLG is dissolved in acetone or acetonitrile at 5–20 mg/mL and introduced into an aqueous stabilizer solution under magnetic or high-shear antisolvent precipitation. Poloxamer 188 or D-α-tocopheryl polyethylene glycol succinate at 0.1–0.5% w/v yields nanoparticles with z-average diameter 100–300 nm and polydispersity index below 0.2 by dynamic light scattering per ISO 22412:2017. Paclitaxel or docetaxel loading is performed by co-dissolution at 5–10% w/w relative to polymer, with encapsulation efficiency above 70% under low aqueous pH or with cationic lipids. Residual acetone or acetonitrile is removed by rotary evaporation at <30 °C, and the dispersion is sterile-filtered through 0.22 µm PVDF. Lyophilization with 5% w/v mannitol or trehalose as cryoprotectant produces a cake with residual moisture <1.0% w/w. Sterility assurance follows ISO 13485:2016 process controls and USP <71>. In vitro cytotoxicity is assessed by ISO 10993-5 extraction method using L929 fibroblasts; endotoxin limits per USP <85> are typically below 0.5 EU/mL for intravenous infusion. The 65:35 ratio provides a mid-range degradation rate that balances drug release over 2–6 weeks in physiological medium, while acid-terminal functionality facilitates surface charge regulation for lysozyme or siRNA complexation. Production-scale homogenization at 10,000–15,000 psi through microfluidizers can reduce batch-to-batch size variation but induces shear-mediated molecular weight loss if the polymer is processed above 20 °C for more than 10 min.

In-Situ Depot Formation and Solvent Exchange Kinetics

In-situ gelling formulations combine 20–35% w/w 65:35 DL-PLG with N-methyl-2-pyrrolidone or dimethyl sulfoxide; upon contact with aqueous tissue fluid, the water-miscible solvent exchanges with water and precipitates the polymer as a monolithic depot. Syringeability through a 21 G needle requires solution viscosity below approximately 800 cP at 25 °C, which places an upper limit on polymer concentration and requires an acid-terminated grade with low to mid inherent viscosity. Solvent exchange rate controls the initial burst: formulations with 30% w/w polymer can release 20–40% of a hydrophilic peptide within 2 h if a highly water-miscible solvent is used. Less aggressive solvents with higher octanol-water partition coefficients reduce early water penetration but prolong the syringeability window only if viscosity remains acceptable. Terminal finished configurations include periodontal and subcutaneous depots, where the polymer phase is packaged in a two-syringe mixing system and combined with the peptide solution immediately before administration. Residual N-methyl-2-pyrrolidone is controlled under ICH Q3C class 2 limits; in vivo tolerability is evaluated by ISO 10993-6 intramuscular implantation in rabbits or rats. Gamma irradiation of the polymer-solvent system at 25 kGy can produce free radicals from the solvent and reduce polymer molecular weight; sterile filtration of the solvent phase and aseptic filling are therefore preferred for terminal sterilisation. Batch release includes viscosity, density, pH of reconstituted formulation, and gelation time in phosphate-buffered saline at 37 °C, which should fall between 10–100 s for a subcutaneously injectable depot.

Thermoplastic Extrusion of 65:35 DL-PLG Rods Requires a Tight 40–48 °C Processing Window

Hot-melt extrusion of 65:35 DL-PLG into bioresorbable drug-eluting rods is performed on a co-rotating twin-screw extruder with L/D 25:1 to 40:1, barrel temperatures from 35 °C to 55 °C, and die temperature held at 40–48 °C. The amorphous copolymer softens rather than melts, so viscous dissipation from screw speeds above 150 rpm can generate localized temperatures exceeding 70 °C and initiate chain scission. Pre-drying under vacuum at 30 °C for 24 h below <0.1% w/w residual moisture is required because hydrolytic degradation during extrusion reduces inherent viscosity and widens die swell. Active pharmaceutical ingredients are blended at 5–30% w/w with gentle tumble or low-shear mixing before extrusion; higher drug loads of >30% w/w depress melt viscosity and produce strand breakage unless die pressure is maintained below 20 bar. The extrudate is drawn through a water-free air cooling system to avoid stress cracking, then cut into rods or pellets for terminal packaging. E-beam or gamma irradiation at 10–25 kGy according to ISO 11137-1 lowers molecular weight, so irradiation dose mapping and post-irradiation inherent viscosity testing are used to confirm that release duration remains within specification. Mechanical testing of compression-molded reference samples follows ASTM D638-14 Type V dimensions; typical tensile strength for unfilled 65:35 DL-PLG films is 30–50 MPa, but the rods are not structural load-bearing devices. Residual monomers are measured by gas chromatography with flame ionization detection, and total residual lactide/glycolide is maintained below 2.0% w/w on the manufacturer’s certificate of analysis. Lot-specific inherent viscosity and residual solvent data must be consulted before setting barrel temperature profiles, because minor lot-to-lot differences in acid-terminal content and moisture uptake shift the effective processing window.

Casting 65:35 DL-PLG films from a ternary solvent system of chloroform, acetone, and 5–15% w/w triethyl citrate produces flexible bioresorbable membranes with dry thickness 50–300 µm. The polymer is dissolved at 5–15% w/w in chloroform, cast onto a release liner via knife-over-roll with gap set at 200–600 µm, and dried in a solvent-retaining chamber with controlled nitrogen flow to prevent condensation. Residual chloroform is removed below ICH Q3C limits by vacuum drying at 35 °C for 48 h. The dried film is die-cut into patches or intrauterine inserts; mechanical tensile properties per ASTM D882-18 typically show ultimate tensile strength 10–40 MPa, elastic modulus 0.5–2.0 GPa, and elongation at break 3–10% depending on plasticizer content. In phosphate-buffered saline at 37 °C, the 65:35 film absorbs water, undergoes bulk degradation, and loses mechanical integrity within 2–6 weeks, which makes it suitable for temporary barrier or drug-loaded mucosal applications but not for orthopaedic load-bearing devices. Ethylene oxide sterilization is performed at 37–45 °C with 24 h aeration; gamma irradiation is avoided when plasticized films are sensitive to crosslinking or free-radical degradation. Residual ethylene oxide, ethylene chlorohydrin, and ethylene glycol limits are verified per ISO 10993-7:2008. Because the acid-terminated polymer hydrolyzes more rapidly in the plasticized matrix than in solvent-free extrusions, accelerated aging at 40 °C/75% RH is used only after confirming that the film remains below its glass transition during storage.

When Electrospinning Is Deployed for Soft Tissue Conduits

When electrospinning is performed with 65:35 DL-PLG for soft tissue conduits, the polymer is dissolved at 10–25% w/w in 3:1 v/v dichloromethane/dimethylformamide or in hexafluoro-2-propanol to obtain a stable jet. A syringe pump delivers the solution at 0.5–3.0 mL/h through a 22 G needle charged to 12–25 kV positive potential, with a grounded rotating drum collector positioned 10–20 cm from the tip. Relative humidity above 40% causes phase separation and bead defects; the chamber is therefore purged with dry nitrogen and maintained at 20–30 °C. Fiber diameter is controlled by solvent ratio, concentration, and flow rate, typically yielding 300–1200 nm for 65:35 DL-PLG. Residual dimethylformamide or hexafluoro-2-propanol is removed by vacuum drying at <30 °C for 24–72 h, and residual solvent content is measured by headspace gas chromatography. Meshes are assessed for suture retention strength and burst pressure where applicable, although published data for this specific 65:35 DL-PLG grade are limited; mechanical anisotropy is quantified by ASTM D638-14 or ring-burst protocols. Cytocompatibility follows ISO 10993-5 extract testing, and subchronic implantation is conducted under ISO 10993-6 in a rodent subcutaneous model. Sterile processing for electrospun conduits usually involves ethylene oxide at 37–45 °C with 12–24 h aeration, because gamma irradiation at 25 kGy reduces fiber tensile strength by approximately 30–60% through chain scission. Terminal products include tubular nerve guides, vascular graft scaffolds, and urological patches where porous ingrowth is required and mechanical load is low.

Particulate leaching with sieved sodium chloride fractions of 150–400 µm is used to fabricate porous 65:35 DL-PLG scaffolds with interconnected void networks. Solution in chloroform at 8–15% w/w is mixed with sodium chloride at 8:1–12:1 porogen-to-polymer mass ratio, cast into PTFE molds, and air-dried for 48 h; leaching in deionized water for 72 h with water changes every 12 h removes salt and creates pores with interconnectivity above 80%. Supercritical CO₂ foaming at 35–40 °C and 80–120 bar with depressurization rates near 1–3 bar/min yields closed-cell morphologies that are less suitable for cell infiltration unless combined with salt leaching. Compressive modulus of scaffolds with 80–90% porosity typically falls below 10 MPa, limiting use to non-load-bearing soft tissue and regenerative niches. In vitro degradation is followed in phosphate-buffered saline at 37 °C over 4–8 weeks, with mass loss, pH drop, and osmolality measured because acid-terminal PLGA accelerates autocatalysis. Ethylene oxide sterilization according to ISO 11135:2014 is preferred over gamma irradiation to preserve the amorphous pore struts. Cytocompatibility is evaluated by ISO 10993-5 direct contact; endotoxin is controlled below 0.5 EU/mL per USP <85>. Terminal finished configurations include cell-carrier discs, periodontal regeneration matrices, and cartilage biopsy scaffolds, where tissue ingrowth rather than long-term load bearing is the primary engineering requirement.