In single-emulsion solvent evaporation, PURASORB PDLG 7504 is dissolved in dichloromethane at 10–25 wt% with a polymer-to-drug mass ratio of 2:1 to 20:1. The active pharmaceutical ingredient is pre-solubilised or pre-suspended in a small aqueous phase and emulsified into the polymer-rich organic phase by a rotor-stator homogeniser operating at 5,000–20,000 rpm. The primary emulsion is transferred into a continuous aqueous polyvinyl alcohol phase at 0.5–2.0 wt%; dichloromethane is removed by overhead agitation at 300–800 rpm and a jacket temperature of 20–35°C over 4–8 h. Solvent removal rate is staged because rapid initial evaporation produces a dense skin that traps residual solvent and raises the final residual dichloromethane burden. The acid-terminated 75:25 copolymer and its inherent viscosity midpoint of 0.4 dL/g are chosen when a depot formulation requires faster water uptake than an ester-capped analogue can provide.

The terminal solid is a lyophilised injectable microsphere powder with particle diameter controlled between 10 µm and 100 µm for intramuscular or subcutaneous administration. Finished-particle diameter is verified by laser diffraction according to USP<429>; injectable formulations are also tested under USP<788> for subvisible particulates and under USP<85> for bacterial endotoxins. Residual dichloromethane is reduced below the ICH Q3C Option 1 permitted daily exposure of 6.0 mg/day by vacuum drying at 25–35°C for 12–48 h. Lyophilisation is performed with 2–5 wt% mannitol or sucrose as a cryoprotectant. Because the acid-terminated 75:25 grade accelerates bulk water uptake, the intraparticle pH can fall during later release phases; co-lyophilisation with magnesium hydroxide at 1–3 wt% is a documented stabilisation approach for acid-labile peptides, but the addition alters burst release and must be characterised for each drug load.

What Limits Burst Release in Antisolvent-Precipitated PLGA Nanoparticles?

When the same acid-terminated 75:25 grade is precipitated from acetone or dimethyl sulfoxide into an aqueous stabiliser solution, the nucleation rate is governed by local polymer supersaturation and interfacial water flux. The organic feed is prepared at 5–20 mg/mL polymer, mixed with an aqueous poloxamer 188 or polyvinyl alcohol phase at an organic-to-aqueous ratio of 1:3 to 1:5, and stirred at 500–1,000 rpm for 10–30 min. The low inherent viscosity midpoint of 0.4 dL/g permits higher organic feed concentration without exceeding a workable membrane flux during purification. The nanoparticle suspension is concentrated by tangential flow filtration through a 100 kDa or 0.1 µm membrane and freeze-dried with 2–5 wt% trehalose or sucrose. Burst release is controlled by the polymer-to-drug ratio, the organic solvent removal rate during precipitation, and residual acetone or dimethyl sulfoxide, both of which are limited under ICH Q3C.

The terminal lyophilised cake is reconstituted into a nanosuspension with Z-average diameter below 300 nm and polydispersity index below 0.2 measured by dynamic light scattering according to ISO 22412:2017. Cytotoxicity is assessed by ISO 10993-5, and sterility of the reconstituted product is established by terminal 0.22 µm membrane filtration before lyophilisation. Published data for this specific grade in continuous antisolvent precipitation is limited; batch-to-batch particle-size variation is therefore expected when moving from laboratory rotor-stator units to production microreactors. The acid terminal groups improve colloidal stability in the aqueous receiving phase, but they also increase the negative zeta potential in the presence of residual polyvinyl alcohol; this interaction must be monitored by electrophoretic light scattering because it affects redispersibility after lyophilisation.

An acid-terminated PLGA of 75:25 molar ratio and inherent viscosity midpoint 0.4 dL/g is also used in N-methyl-2-pyrrolidone-based in-situ forming depots without forming discrete particles. The polymer and drug are co-dissolved in NMP at 35–50 wt% polymer, producing a solution with dynamic viscosity between 200 mPa·s and 2,000 mPa·s at 25°C measured by rotational rheometry according to ISO 3219. The solution is filled into a pre-filled syringe and administered through an 18 G to 21 G needle. Injection-force testing on a texture analyser at 100 mm/min shows that the upper practical concentration for this grade is governed by glide force rather than thermodynamic solubility; published data specific to PDLG 7504 is limited above 50 wt% polymer in NMP, where the solution no longer passes through the designated needle with acceptable manual injection force.

Upon contact with aqueous tissue fluid, NMP diffuses outward and water diffuses inward, precipitating the acid-terminated 75:25 polyester into a depot that conforms to the injection site. The terminal product is a subcutaneous or intratumoral implant generated in situ, with release controlled by diffusion through the precipitated polymer-rich shell and subsequent bulk erosion. Residual NMP is controlled under ICH Q3C Option 1 PDE of 5.3 mg/day. The formulated syringe is tested under ISO 10993-6 for local tissue reaction after implantation and under USP<85> for bacterial endotoxins. This route is not suitable for drugs with poor stability in NMP or for dose volumes below 0.2 mL, because the depot geometry becomes poorly reproducible when the injected volume cannot maintain a coherent mass.

Hot-Melt Extrusion Through an L/D 30:1 Co-Rotating Twin-Screw Line

Thermal processing of PDLG 7504 for implant rods is performed on a co-rotating twin-screw extruder with a length-to-diameter ratio of 30:1 to 40:1, segmented screws, and a gravimetric feeder. The acid-terminated 75:25 grade is pre-dried at 25–35°C under vacuum to a Karl Fischer moisture content below 0.1 wt%; barrel set points are established between 90°C and 120°C, which is above the glass transition range of 45–55°C measured by differential scanning calorimetry according to ISO 11357-2. Screw speed is maintained between 20 rpm and 100 rpm, and melt residence time is held below 5 min to limit autocatalytic hydrolysis and transesterification. The extrudate is pulled through a cooling air knife, cut into cylindrical rods of 1.8 mm to 3.5 mm diameter, and packaged under nitrogen.

The terminal implant is tested for weight uniformity, burst release, and mechanical integrity. Because the D,L-lactide-rich composition is amorphous, the rods exhibit plastic deformation rather than brittle failure; the 75:25 ratio gives a degradation profile that is slower than 50:50 PLGA but faster than 85:15 PLGA. Residual moisture is monitored by Karl Fischer titration according to USP<921>, and release testing is conducted in USP<711> apparatus 4 or apparatus 7 for extended-release implants. ISO 10993-6 implantation testing is required for the final device, and gamma irradiation at 25 kGy must be preceded by moisture and oxygen control because free-radical chain scission reduces molecular weight and shifts the release lag phase; published data for this specific grade after terminal sterilisation at 25 kGy is limited.

When Electrospinning Turns Acid-Terminated PLGA into Fibrous Drug-Eluting Meshes

Electrospinning of PDLG 7504 at 10–25 wt% in dichloromethane–dimethylformamide mixtures of 3:1 to 1:1 v/v produces nonwoven fibre mats with diameters from 200 nm to 5 µm, depending on applied voltage, feed rate, and collector distance. A positive displacement syringe pump delivers the solution at 0.5–2.0 mL/h while a high-voltage supply applies 10–20 kV to a metallic needle positioned 10–20 cm from a grounded rotating drum or static plate. Relative humidity is held below 40% to avoid premature fibre fusion and water-driven hydrolysis. The 75:25 D,L-lactide-rich copolymer remains amorphous in the fibre, which reduces the Young’s modulus of the mesh but improves conformability at surgical sites.

The terminal mesh is cut into implantable patches or rolled into cylindrical depot structures for local drug delivery. Residual dimethylformamide must meet ICH Q3C Option 1 PDE of 8.8 mg/day, and residual dichloromethane must meet the 6.0 mg/day PDE. Fibre diameter distribution is determined by scanning electron microscopy, and in vitro release is assessed by USP<711> dissolution testing. ISO 10993-5 cytotoxicity and ISO 10993-6 implantation studies apply to the finished device. Published data for this specific grade in electrospun format is limited with respect to long-term dimensional stability, because amorphous D,L-lactide segments allow stress relaxation and bulk erosion that can reduce fibre integrity after 4–6 weeks of aqueous exposure.

Spray coating with low-viscosity PLGA solutions avoids the melt residence time problem altogether but replaces it with a more demanding drying and residual solvent control sequence. PDLG 7504 is dissolved at 2–10 wt% in ethyl acetate or acetone and deposited through an ultrasonic nozzle operating at 20–60 kHz onto rotating fixtures carrying cardiovascular or orthopaedic device surfaces. The solution is applied in 10–30 passes with inter-pass drying at 35–45°C for 30–90 s per pass, because thick wet layers trap solvent and create blister defects when the outer film forms before the inner solvent has diffused to the surface.

The terminal component is a drug-eluting implant or coated device with a coating thickness between 5 µm and 50 µm measured by profilometry or scanning electron microscopy. Residual ethyl acetate and acetone are controlled under ICH Q3C Class 3 limits of 50 mg/day; no Class 2 solvent is introduced unless the formulation requires a co-solvent such as dimethylformamide. The coated device is tested under ISO 10993-5 for cytotoxicity and ISO 10993-6 for local tissue reaction. Adhesion is evaluated by a tape test adapted from ISO 2409, and coating delamination at the polymer–substrate interface is the principal failure mode when the substrate has not been plasma-cleaned before coating.