For sustained-release injectable microspheres, Resomer RG 504 H is processed by a double-emulsion solvent evaporation route when the API is a water-soluble peptide or by a solid-in-oil-in-water dispersion when the API is a poorly water-soluble small molecule. The dispersed phase is commonly formulated at 20–30% w/v RG 504 H in dichloromethane or ethyl acetate, with dichloromethane preferred for higher encapsulation of low molecular weight peptides but controlled below the ICH Q3C(R8) Option 2 limit of 600 ppm in the finished microsphere. The aqueous continuous phase contains 0.5–1.0% w/v poly(vinyl alcohol) with 87–89% hydrolysis degree. Rotor-stator homogenization at 10,000–15,000 rpm for 2–5 min generates the primary emulsion, and the secondary water-in-oil-in-water dispersion is stirred at 400–700 rpm under relative humidity below 40% to limit premature PLGA hydrolysis. Solvent removal is conducted by vacuum distillation at 300–400 mbar and 25–35°C for up to 4 h, followed by hardening with ice-cold aqueous washing. The microspheres are lyophilized with mannitol or trehalose at 2–5% w/v as cryoprotectant. The terminal dosage form is a sterile lyophilized cake for reconstitution with Water for Injection, tested for release using USP 711 Apparatus 4 at 37 ± 0.5°C in phosphate-buffered saline pH 7.4 containing 0.02% sodium azide, and checked for particulate matter according to USP 788. In the acid-terminated 50:50 PLGA matrix, bulk hydrolysis releases lactic and glycolic acid monomers that lower the intraparticle pH below 4.0; this acidic microclimate can stabilize pH-dependent peptides but limits use with acid-labile actives. Production-scale transfer failures include bimodal particle size distributions when rotor-stator tip speed is not held constant and residual dichloromethane increase when ethyl acetate replacement is attempted without adjusting the vacuum ramp rate.

What Governs Burst Release and Depot Viscosity During Phase Inversion of an Acid-Terminated PLGA/NMP Solution?

In situ forming implant depots based on RG 504 H are usually formulated as sterile injectable solutions in N-methyl-2-pyrrolidone at polymer concentrations of 30–40 wt%, with optional addition of 5–15 wt% triacetin or propylene carbonate to modulate solvent efflux and depot porosity. Upon injection into aqueous subcutaneous tissue, the water-miscible solvent exchanges with tissue fluid and induces phase inversion, forming a porous degradable depot. The acid-terminated 50:50 PLGA grade absorbs water and enters the release lag phase faster than ester-terminated grades; therefore burst release is controlled by increasing polymer concentration, reducing NMP fraction, and selecting a 20–23 G needle to avoid injection force variations. Depot viscosity below 2.0 Pa·s at 25°C is generally required for manual injection through a 21 G needle, placing an upper practical polymer concentration near 35–40 wt% for this grade. In vitro release is measured with USP 711 Apparatus 4 or a shaking-bath system using 40 mL phosphate buffer at 37°C. Because N-methyl-2-pyrrolidone is a Class 2 residual solvent under ICH Q3C(R8), local tolerance is evaluated according to ISO 10993-6 and the solvent burden must be justified by toxicological data. Gamma irradiation at 25 kGy can reduce PLGA molecular weight and shift release earlier; if terminal sterilization is required, dose-mapping under ISO 11137-1:2006/Amd 1:2013 is applied. Sterile filtration is generally not feasible above 30 wt% because solution viscosity exceeds 0.45 µm membrane capacity, so aseptic manufacturing or terminal sterilisation is the primary route. The terminal product is a single prefilled syringe forming a weekly-to-monthly depot.

When intravenous compatibility or high tissue permeability is required, nano-precipitation of Resomer RG 504 H is applied to hydrophobic API payloads. A typical organic phase contains 1.0–5.0 mg/mL RG 504 H and 0.1–0.5 mg/mL API in acetone, acetonitrile, or a 70:30 v/v acetone-ethanol mixture; the aqueous anti-solvent phase contains 0.05–0.2% w/v polysorbate 80 or poloxamer 188 as colloidal stabilizer. The organic phase is injected at 5–10 mL/min into the aqueous phase under overhead stirring at 500–1000 rpm, maintaining an organic-to-aqueous ratio of 1:5 to 1:20. Diffusion-controlled nucleation and growth produce particles with a z-average diameter of 80–200 nm and polydispersity index 0.10–0.20 by dynamic light scattering according to ISO 22412:2017. Aggregates are removed by centrifugation at 10,000 × g for 10 min or tangential flow filtration through a 100 kDa polyethersulfone membrane. Residual organic solvent is reduced by rotary evaporation at 25–30°C and 200–300 mbar to below ICH Q3C(R8) limits, such as 5,000 ppm for acetone. The terminal product is a lyophilized nanocrystalline suspension for intravenous infusion after reconstitution with isotonic saline containing 2–5% w/v trehalose. For active substances with log P greater than 3, drug loading in the PLGA nanoparticle typically remains below 10% w/w without precipitation inhibitors. Hydrophobic ion pairing can increase loading but introduces anionic surfactant compatibility issues in USP 788 and ISO 10993-5 tests. Cationic stabilizers are generally avoided because the acid-terminated PLGA forms electrostatic aggregates that increase particle size during storage.

Hot-Melt Extruded Rod Implants and Near-Net-Shape Compression Moulding Conditions

Implantable rods and wafers are produced from RG 504 H by hot-melt extrusion or compression moulding after pre-drying at 25–30°C under vacuum below 20 mbar for at least 12 h until moisture content is below 0.1% w/w by Karl Fischer titration. The glass transition temperature of the acid-terminated 50:50 PLGA grade is 44–48°C by differential scanning calorimetry according to ASTM D3418-15, but hot-melt extrusion requires barrel temperatures of 120–150°C to lower melt viscosity sufficiently. A co-rotating twin-screw extruder with 16 mm screw diameter and L/D 40:1 is operated at screw speeds of 100–200 rpm, with zone temperatures of 110°C, 130°C, and 140°C at the die, and melt pressure between 20–60 bar. The processing window is narrow because residence time above 2 min or melt temperature above 160°C accelerates hydrolytic and thermomechanical chain scission, causing a measurable reduction from the specified inherent viscosity range of 0.45–0.60 dL/g. A torque increase of more than 20% from baseline indicates degradation and requires purging with fresh polymer. The extruded strand is quenched in a dry nitrogen stream, cut into cylindrical rods of 1.0–2.0 mm diameter and 2–3 cm length, and heat-sealed in aluminium foil pouches with desiccant. Compression moulding for wafers uses 200–400 mg of polymer-drug blend in a 10 mm diameter die pressed at 70–80°C under 5–10 kN for 2–5 min. The terminal products are surgically placed subcutaneous rods or intra-tumour wafers; in vitro release is tested with USP 711 Apparatus 4 in 50 mL phosphate buffer at 37°C, and mechanical integrity is followed by diameter stability and fracture resistance during the release interval. Pre-drying at relative humidity above 60% is mandatory to prevent bubbles, torque drift, and batch-to-batch inherent viscosity scatter on production extruders.

As a downstream intermediate for poorly water-soluble APIs intended for oral solid dosage forms, spray-dried amorphous solid dispersions of Resomer RG 504 H are used when microsphere or implant formats are not required. The feed solution is prepared at 2–5% w/v total solids in acetone or dichloromethane with a drug-to-polymer ratio between 10:90 and 30:70 by weight. Spray drying on a Büchi B-290 laboratory unit uses a two-fluid nozzle at atomizing gas pressure 1.0–1.5 bar, inlet temperature 55–65°C, outlet temperature 25–40°C, and feed rate 5–10 mL/min. The collected powder is secondary dried under vacuum at 30°C for 6–12 h to reach residual acetone below 5,000 ppm or residual dichloromethane below 600 ppm according to ICH Q3C(R8). X-ray powder diffraction per USP 941 confirms absence of API crystallinity, and modulated differential scanning calorimetry shows a single glass transition above 45°C without low-temperature melting endotherms. Compression into tablets requires dry granulation because the low glass transition and hygroscopic nature of the spray-dried dispersion cause sticking on rotary tablet press punches at ambient relative humidity above 55%. The final dosage form is a conventional immediate-release or delayed-release tablet; dissolution is tested by USP 711 Apparatus II at 50–75 rpm in simulated gastric fluid or phosphate buffer with 0.1–0.5% sodium lauryl sulfate. Acid-terminated PLGA undergoes ester hydrolysis in acidic media, which can introduce dissolution testing artefacts if the medium is not buffered or if samples are held for more than 24 h before analysis.

If Gamma Irradiation Is Specified for Bioresorbable Drug-Eluting Coatings on Metallic Stents

On metallic stents requiring a low-thickness bioresorbable drug-eluting coating, Resomer RG 504 H is applied as the drug-carrying matrix when a controllable release layer is needed without permanent polymer residue. The coating solution is prepared at 0.5–2.0% w/v total solids in a 70:30 v/v acetone-methanol mixture, with the API dissolved or suspended at 10–25% w/w of polymer weight. Ultrasonic spray coating is performed with a nozzle frequency of 60–120 kHz, a flow rate of 0.1–0.5 mL/min, and a nozzle-to-substrate distance of 8–15 mm; the rotating mandrel is held at 25–35°C under a dry nitrogen purge. Coating thickness is built in multiple passes to 2–10 µm, followed by vacuum annealing at 40–50°C for 12–24 h to remove residual solvent. If terminal sterilisation by gamma irradiation is required, the absorbed dose is typically 25 kGy under ISO 11137-1:2006/Amd 1:2013, with dose-mapping across the packaged stent to confirm minimum and maximum dose. Irradiation causes chain scission in the acid-terminated 50:50 PLGA, shifting release earlier and reducing coating toughness; therefore pre-irradiation stability studies must define the compensating increase in API load or coating thickness. Compatibility is evaluated by ISO 10993-4 haemocompatibility, ISO 10993-5 cytotoxicity, and USP 788 particulate matter after simulated deployment. The main operational boundary is hydrolytic instability: the coating solution should be prepared immediately before spraying because the acid-terminated PLGA autocatalytically degrades in acetone-methanol if stored at room temperature for more than 8 h. The terminal product is a crimped stent sealed in a sterile barrier pouch per ISO 11607-1:2019.