Comparison of PLA and PLGA in Drug Delivery

Research will be limited to:
- Material: PLGA copolymer
- Application: Drug Delivery
- Property: Drug release kinetics
- Parameter: PLGA copolymer ratio

Strategy: To provide a precise recommendation for selecting the PLGA copolymer ratio, I need to establish the quantitative relationship between the lactic acid (LA) to glycolic acid (GA) ratio and the resulting degradation rates and drug release profiles. Specifically, I will look for comparative studies that map 50:50, 65:35, 75:25, and 85:15 ratios against release kinetics for different drug classes (hydrophilic vs. hydrophobic). I will also investigate the impact of the ratio on the glass transition temperature (Tg) and crystallinity, as these dictate the diffusion-controlled vs. erosion-controlled phases of release.
Queries:
- PLGA lactic glycolic ratio drug release kinetics correlation
- PLGA 50:50 vs 75:25 degradation rate drug delivery comparison
- effect of PLGA copolymer ratio on burst release and lag phase
- PLGA copolymer ratio influence on glass transition temperature and drug diffusion
- mathematical modeling PLGA ratio drug release kinetics

Executing 5 concurrent searches:
- PLGA lactic glycolic ratio drug release kinetics correlation
- PLGA 50:50 vs 75:25 degradation rate drug delivery comparison
- effect of PLGA copolymer ratio on burst release and lag phase
- PLGA copolymer ratio influence on glass transition temperature and drug diffusion
- mathematical modeling PLGA ratio drug release kinetics

Found 8 papers. Key excerpts:

Ibuprofen-loaded poly(lactic-co-glycolic acid) films for controlled drug release — Jianmei Pang et al., 2022

IntroductionNowadays, poly(lactic-co-glycolic acid) (PLGA) is widely applied in controlled drug delivery systems due to its biodegradability, toxicological safety, and good biocompatibility.Citation1PeruginiPGentaIContiBPeriodontal delivery of ipriflavone: new chitosan/PLGA film delivery system for a lipophilic drugInt J Pharm20032521912550776 [Crossref], [Web of Science ®], [Google Scholar]–Citation5KloseDSiepmannFWillartJFDrug release from PLGA-based microparticles: effects of the “microparticle:bulk fluid” ratioInt J Pharm201038312313119748558 [Crossref], [Web of Science ®], [Google Scholar] Specifically, PLGA-based carriers have been used in long-term drug delivery systems because they have the potential to control drug release from a few days up to several months.Citation6XuQCzernuszkaJTControlled release of amoxicillin from hydroxyapatite-coated poly (lactic-co-glycolic acid) microspheresJ Contr Release2008127146153 [Crossref], [Web of Science ®], [Google Scholar],Citation7Fernández-CarballidoAHerrero-VanrellRMolina-MartinezITBiodegradable ibuprofen-loaded PLGA microspheres for intraarticular administration effect of Labrafil addition on release in vitroInt J Pharm2004279334115234792 [Crossref], [Web of Science ®], [Google Scholar] In general, PLGA can degrade into water soluble, non-toxic products of normal metabolism through hydrolysis which is important for its practical application.Citation8SchlieckerGSchmidtCFuchsSHydrolytic degradation of poly(lactide-co-glycolide) films: effect of oligomers on degradation rate and crystallinityInt J Pharm2003266394914559392 [Crossref], [Web of Science ®], [Google Scholar] Moreover, PLGA is one of the few synthetic polymers which have been approved for human clinical use. Several products such as Lupron Depot based on PLGA microparticles are

available on the market.Citation5KloseDSiepmannFWillartJFDrug release from PLGA-based microparticles: effects of the “microparticle:bulk fluid” ratioInt J Pharm201038312313119748558 [Crossref], [Web of Science ®], [Google Scholar],Citation9WangJWangBMSchwendemanSPCharacterization of the initial burst release of a model peptide from poly(D,L-lactide-co-glycolide) microspheresJ Contr Release200282289307 [Crossref], [Web of Science ®], [Google Scholar] However, PLGA-based carriers are mostly solid particles including nanoparticles, microparticles, and microcapsules, which have the disadvantages of low drug loading capacity, easy aggregation, and polydispersable particle sizes.Citation10AcharyaGShinCSVedanthamKA study of drug release from homogeneous PLGA microstructuresJ Contr Release2010146201206 [Crossref], [Web of Science ®], [Google Scholar] Therefore, it is necessary to prepare aspherical carriers instead of particles. Various biocompatible and biodegradable polymer films have been widely used for controlled drug release due to their excellent characteristics, such as high surface area, softness, absorbency, and ease of fabrication into many product forms.Citation11JacksonJKSmithJLetchfordKCharacterization of perivascular poly (lactic-co-glycolic acid) films containing paclitaxelInt J Pharm20042839710915363506 [Crossref], [Web of Science ®], [Google Scholar]–Citation16ShiSAWangXHGuoGFanMHuangMJQianZYPreparation and characterization of microporous poly(D,L-lactic acid) film for tissue engineering scaffoldInt J Nanomed2010510491055 [Taylor & Francis Online], [Web of Science ®], [Google Scholar] Until now, a number of methods have been developed for fabrication of polymer films, such as layer-by-layer (LBL) methodCitation17JiangBBLiBYTunable drug loading and release from polypeptide multilayer nanofilmsInt J Nan

omed200943753 [Taylor & Francis Online], [Web of Science ®], [Google Scholar] and solvent evaporation technique.Citation18PatelNAPatelNJPatelRPDesign and evaluation of transdermal drug delivery system for curcumin as an anti-inflammatory drugDrug Dev Ind Pharm20093523424218785045 [Taylor & Francis Online], [Web of Science ®], [Google Scholar] However, the reported methods usually involve some cumbersome processing steps and are time-consuming. Therefore developing a simple route to prepare drug-loaded fibers is still a great challenge.Ibuprofen (IBU) is classified by the Biopharmaceutics Classification System (BCS) as a Class II active pharmaceutical ingredient as it presents low solubility and high permeability.Citation19MoneghiniMBellichBBaxaPMicrowave generated solid dispersions containing IbuprofenInt J Pharm200836112518573321 [Crossref], [Web of Science ®], [Google Scholar] Thus, increasing the solubility of IBU to enhance its bioavailability is the major obstacle.Citation20PangJLuanYLiFIonic liquid-assisted synthesis of silica particles and their application in drug releaseMater Lett20106425092512 [Crossref], [Web of Science ®], [Google Scholar],Citation21NewaMBhandariKHLiDXPreparation, characterization and in vivo evaluation of ibuprofen binary solid dispersions with poloxamer 188Int J Pharm200734322823717597315 [Crossref], [Web of Science ®], [Google Scholar] Previous studies indicate that the formulated IBU microspheres have a high initial burst resulting from the accumulation of drug crystals on the surface of microspheres.Citation8SchlieckerGSchmidtCFuchsSHydrolytic degradation of poly(lactide-co-glycolide) films: effect of oligomers on degradation rate and crystallinityInt J Pharm2003266394914559392 [Crossref], [Web of Science ®], [Google Scholar],Citation22JainJAThe manufacturing techniques of various drug loaded biodegradable poly (lactide-co-glycolide) (PLGA) devicesBiomaterials2000212475249011055

3D printed PLGA implants: How the filling density affects drug release — C. Bassand et al., 2023

Poly (D,L lactic- co -glycolic acid) (PLGA)-based drug delivery systems offer an interesting potential for controlled drug delivery applications [[1], [2], [3], [4], [5]] because they: (i) are completely biodegradable, (ii) are biocompatible [6], and (iii) can provide flexible release rates and periods [[7], [8], [9], [10]]. A large variety of drugs has been incorporated into this type of drug delivery systems, including small and large molecules as well as freely and poorly water soluble compounds [[11], [12], [13]]. Different types of systems have been proposed and reached the market [[14], [15]]: Pre-formed implants and microparticles as well as liquid formulations, which are injected into the patient's body and harden in vivo (“in-situ forming implants”).
To adjust desired release kinetics for a given drug, various formulation and processing parameters can be varied. This includes for instance the drug content, type of PLGA (differing in the average polymer molecular weight, type of end groups and lactic acid: glycolic acid ratio), geometry and dimensions of the device, addition of further excipients as well as the manufacturing procedure. The impact of the latter is not to be underestimated, because it can strongly affect the resulting inner system structure, which can be of crucial importance for the conditions of drug release. For example, PLGA-based microparticles can be prepared via different solvent extraction/evaporation methods, using oil-in-water (O/W) emulsions or water-in-oil-in-water (W/O/W) emulsions [[16], [17]]. In the first case, non-porous microparticles with relatively slow drug release can be obtained, whereas in the latter case highly porous systems can be prepared with much higher drug release rates (even if the qualitative and quantitative compositions of the microparticles are very similar): The presence of numerous inner pores can facilitate drug transport within the microparticles.
Despite the considerable practical importance of PLGA-based controlled drug delivery systems, the physico-chemical phenomena which are involved in the control of drug release are often not fully understood. This hampers the development of new drug products: Cost-intensive and time-consuming series of trial-and

-error experiments are required, and unexpected tendencies can be observed. This can in great part be attributed to the complexity of the underlying drug release mechanisms [[18], [19], [20], [21], [22], [23]]. Upon contact with aqueous media, water penetrates into the systems and rather rapidly wets the entire devices. This results in ester bond cleavage throughout the systems (“bulk degradation”) [24]. If the drug is present in the form of particles (crystalline or amorphous), the latter dissolve in the water and the drug becomes mobile. Due to the concentration gradients, the dissolved drug subsequently diffuses out of the system. Depending on the initial loading and solubility of the drug in the wetted PLGA matrix, dissolved and non-dissolved drug might co-exist during considerable periods of time. Importantly, only dissolved drug is mobile and can diffuse out. This diffusional mass transport might occur through water-filled pores or channels and/or through an intact, more or less swollen polymeric network. The affinity of the drug to the PLGA and to water, as well as its molecular weight affect the importance of these different diffusion pathways. Certain drugs also act as plasticizers for PLGA, increasing the mobility of the macromolecular chains [25]. Furthermore, the hydrolytic cleavage of each ester bond generates a new -COOH end group. So, the pH within the drug delivery system might locally substantially drop [[26], [27], [28], [29]], leading to autocatalytic effects: Ester bond cleavage is catalyzed by protons. Depending on the dimensions of the drug delivery systems, water-soluble generated acids can more or less rapidly diffuse out into the surrounding environment, so that the micro pH more or less drastically decreases. Accelerated polymer degradation leads to increased porosity and increased drug mobility. Certain systems have been reported to become highly porous at their center over time [16]. These dynamic changes in the system's structure upon exposure to the release medium can strongly affect drug mobility. In addition, polymer swelling has been suggested to play a key role for the control of drug release from PLGA-based implants and microparticles [[30], [31]].
Different manufacturing techniques can be applied to prepare PLGA implants for controlled drug delivery, including for instance compression [32], casting [33], hot

melt extrusion [34] and 3D printing [[35], [36]]. The latter process offers the advantage to allow for personalized medication [[37], [38], [39], [40], [41]]: Depending on the specific needs of each patient (e.g. dosage, release rate, release period, combination of drugs), personalized implants can be printed. In addition, the implant's shape and geometry can be adapted to each patient. However, yet relatively little knowledge is available on the processability of PLGA-drug blends, the resulting drug release kinetics and the possibilities to adjust desired drug release kinetics. One of the simplest parameters which can be varied during implant printing is the “filling density” of the system: The implant can be designed as a non-porous device (100% filling density), or as a highly porous system (e.g. 10% filling density). However, so far it is unclear how drastically the filling density of the implant impacts the resulting drug release patterns.
The aim of this study was to better understand the importance of the filling density of 3D printed PLGA implants for the control of drug release. Ibuprofen, a small, acidic molecule with anti-inflammatory activity, was studied as the drug. An Arburg Plastic Freeforming (APF) printer was used to prepare different types of implants, which were thoroughly characterized before and after exposure to different release media in two experimental set-ups. The classical exposure to well agitated bulk fluid (phosphate buffer pH 7.4) was compared with the exposure to 0.5% agarose gels: aiming to better mimic the presence of living tissue surrounding the implant in the patient's body [[42], [43], [44]]. In addition, gravimetric measurements, optical microscopy, Differential Scanning Calorimetry, Gel Permeation Chromatography, and Scanning Electron Microscopy were applied to better understand the underlying mass transport mechanisms in the different types of implants under the investigated conditions.

The Effect of Polymer Blends on the In Vitro Release/Degradation and Pharmacokinetics of Moxidectin-Loaded PLGA Microspheres. — Hongjuan Zhang et al., 2023

Poly (D, L lactic-co-glycolic acid) (PLGA), certified by the FDA, has been employed extensively for sustained-release drug delivery in the past few decades due to its attractive biocompatibility and biodegradable characteristics [ 1 , 2 , 3 ]. PLGA microspheres are one of the best studied long-acting drug delivery systems, as they can encapsulate a variety of drugs and the drug release rate can easily be tailored by regulating the polymer’s characteristics, such as molecular weight, terminal group and the ratio of two monomers [ 4 , 5 , 6 , 7 ]. PLGA microspheres display a variety of benefits over conventional drug delivery methods, including prolonged drug release duration, decreased dosing frequency, improved patient compliance, reduced toxicity and low medical costs [ 8 ]. There are several commercially available PLGA microspheres on the market, including Vivitrol ® , Risperidal ® , Lupron Depot ® , Bydureon ® and Perseris™ [ 9 , 10 , 11 , 12 , 13 ]. Therefore, it is promising to develop PLGA microspheres for loading various therapeutic agents to achieve long-term drug delivery.
The desired long-acting injections should release the encapsulated drugs in nearly zero-order within the required time period. However, the in vitro release of drugs from PLGA microspheres typically involves three phases, including a burst release, followed by a long lag phase, and finally a near zero-order fast release [ 14 , 15 ]. The long lag phase during which only a very small amount of drug is released from microspheres is an important problem in the development of microspheres, as it could lead to insufficient blood drug concentration for treatment [ 16 ]. In order to eliminate the long lag phase, additives, for example porogens, are usually added to the formulation to modify release. However, the addition of additives may also result in local irritation and toxicity [ 17 ]. Hence, using polymer blending to prepare microspheres is a good strategy for adjusting drug release.
Many studies have reported using polymer blends to modify the release of microspheres [ 6 , 17 , 18 , 19 , 20 ]. Leuprolide acetate microspheres were prepared using PLGA polymer blends to reduce the burst release of microspheres [ 6 ]. By blending different PLGAs in various ratios, ganciclovir microspheres were prepared to effectively change the

drug release rate of microspheres [ 17 ]. Accordingly, the lag phase caused by the slow degradation of high-molecular-weight PLGA could be minimized or even eliminated by combining high- and low-molecular-weight PLGA. However, there is a lack of a thorough investigation into the practical effects of polymer blends on the in vitro and in vivo performance of hydrophobic small-molecule PLGA microspheres.
In general, drugs are released from PLGA microspheres in two ways: diffusion and degradation/erosion [ 21 , 22 ]. The burst release of microspheres is usually controlled by diffusion, while the lag phase and subsequent fast release are controlled by the erosion of PLGA. There are four recognized processes of drug release from PLGA microspheres, including water absorption and swelling of microspheres, matrix destruction caused by hydrolysis of the polymer chain, internal pore formation and drug diffusion through pores. Therefore, understanding the in vitro release/degradation mechanism is crucial. In addition, a real-time release test at 37 °C is often time-consuming and cumbersome, making it necessary to develop an accelerated release method by changing the release condition.
In this study, we selected moxidectin (MOX), an anthelmintic macrocyclic lactone endectocide [ 23 , 24 ], as the model of a hydrophobic small molecule to evaluate the effect of polymer blends on the in vitro and in vivo performance of hydrophobic small molecule-loaded PLGA microspheres. For this aim, four moxidectin-loaded microspheres (MOX-MS) were prepared using various ratios of polymer blends and characterized for surface morphology, particle size, drug loading and the physical state of the drug within the microspheres. Thereafter, real-time in vitro release and accelerated release tests were performed, as well as an investigation into the in vitro degradation mechanism of microspheres. Furthermore, the pharmacokinetics of MOX-MS were evaluated. These results will provide a sufficient theoretical basis for controlling the release of microspheres by blending polymers so as to eliminate or minimize the lag phase.
MOX-MS with various PLGA blends were prepared using the O/W emulsification solvent evaporation method. No significant differences were observed in the drug loading (36%) and encapsulation efficiency (91%) among four formulations ( Table 1

and Figure 1 ), which may be explained by the high hydrophobicity of moxidectin. Nevertheless, with the increase of low-molecular-weight PLGA, the particle size reduced slightly ( Figure S1 ), which might be attributed to the low viscosity of the organic phase during the preparation process caused by the addition of low-molecular-weight PLGA to the microspheres [ 25 ]. However, these results are contrary to those of previous studies about peptide drugs encapsulated in PLGA blends [ 18 ].
SEM results revealed that the surface of microspheres displayed irregular polygonal wrinkles with a diameter of about 1 μm ( Figure 1 and Figure S2 ). This might be attributed to the high viscosity of the oil phase as well as the high hydrophobicity of the drug and PLGA, leading to the rapid solidification of the surface of the microspheres [ 26 ]. Moreover, four formulations also showed similar morphology and uniform particle size distribution, which was consistent with the results obtained by the laser diffraction particle size analyzer.
Using a powder X-ray diffraction technique, the existing state of moxidectin within the PLGA microspheres was analyzed. As illustrated in Figure 2 , there was no obvious peak of moxidectin and PLGA, indicating that the drug existed in an amorphous state. In addition, there was no significant change in the FT-IR spectra of PLGA, moxidectin and moxidectin-loaded microspheres, indicating that there was no reaction between the drug and PLGA ( Figure S3 ). In summary, these results are beneficial for the release control of microspheres [ 27 ].
The higher the surface hydrophilicity and free energy, the better the compatibility of microspheres. To evaluate the effect of different polymer blends on the surface wettability of microspheres, the water contact angles of different formulations were measured. The water contact angles were 74.13°, 69.40°, 63.20° and 54.70° for F1, F2, F3 and F4, respectively ( Figure 3 ). This indicated that the addition of low-molecular-weight PLGA improved the surface compatibility of microspheres, which may be related to improving the wetting of microspheres and thus accelerating their in vitro release.
Based on the

Effect of copolymer composition on particle morphology and release behavior in vitro using progesterone — Yue Zhang et al., 2018

In the development of oral formulations, the aqueous solubility of the drug plays a key role in the extent of drug absorption [ 1 ]. For any therapeutics to be successfully absorbed in the body, it needs to be present in the form of an aqueous solution at the desired site of action [ 1 , 2 ]. The solubility is the limiting factor for a drug to attain the desired concentration in systemic circulation in order to achieve pharmacological response for an orally administered drug [ 3 ]. However, many drug molecules under development are poorly soluble: that critically limits their absorption, resulting in poor bioavailability and pharmacokinetics in vivo [ 2 , 3 ].
Progesterone is an endogenous steroidal hormone that is involved in all aspects of reproduction and used to control reproductive function. It is being used for hormone replacement therapy (HRT) in menopausal women, who failed to ovulate and stop producing progesterone from their ovaries [ 4 ], this would drop down the hormone levels in the body, hence results in various effects such as hot flashes and night sweats. In addition, recent studies have reported the applicability of progesterone in preventing preterm birth. This is defined by World Health Organization as birth before 37 weeks of pregnancy are completed [ 5 ], which is the largest cause of neonatal morbidity and mortality in infants [ 6 , 7 ]. Progesterone has an essential role in maintaining pregnancy [ 8 , 9 ], and it is involved in establishing uterine quiescence [ 10 ]. The various administration routes of progesterone include oral, parental (intramuscular and subcutaneous) and topical (as a cream of vaginal gel), of which peroral delivery is the most preferred route of administration. However, progesterone suffers from a short half-life and low water solubility; these lead to inconsistent bioavailability and high first pass metabolism [ 11 ]. Biodegradable polymers are promising candidates for oral drug delivery with their ability to shield the drug from external harsh chemical and enzymatic degradation of the gastrointestinal tract (GIT) as well as bypass the first pass metabolism through their unique uptake mechanism [ 12 , 13 ]. A previous study reveals that the dissolution rate of poorly water-soluble drug can be enhanced by incorporating it into a polymeric matrix thereby inducing its amorphous state [ 14 ], again making this a preferable oral delivery technique. In addition to this, long term sustained drug delivery systems can

also be obtained using biodegradable particles, in which progesterone can be administered via depot injection [ 15 ] and released for a prolonged period of time. In a depot injection, the medication is deposited in a localized mass via intramuscularly, subcutaneously or intradermal injection. This offers advantages over long-term daily injection which can compromise patient compliance.
Biodegradable microparticles in the field of controlled drug release, where the drug is continuously released from the particles within a period of time, have generated immense interest. They are able to deliver a wide range of drugs through various administration routes and enhance drug bioavailability and protect the drug from degradation [ 16 ]. The drug release behavior can be controlled by altering polymer compositions including molecular weight, monomers ratio and the nature of terminal groups for copolymers [ 17 ]. It has been reported that other than the physicochemical properties of polymers, factors such as particle size [ 18 , 19 ], morphology and porosity [ 20 ], as well as drug loading [ 21 ] also play important roles in deciding the drug release kinetics. Appropriate selection of the biodegradable polymeric materials is essential to develop a successful particulate delivery system. Among the variety of materials available to fabricate these solid particles, copolymers have received much attention in recent years as different physicochemical properties can be achieved by altering their polymer composition.
Poly ( d , l ‑lactic‑ co ‑glycolic acid) (PLGA) is one of the most frequently used biodegradable polymers in drug delivery applications [ 22 , 23 ]. PLGA is approved by the FDA as a drug delivery vehicle and it preferred for this purpose because of its favorable biodegradability, biocompatibility and tunable physicochemical properties [ 24 , 25 ]. The copolymers with various molecular weight s and copolymer compositions are commercially available. The ability to alter lactide: glycolide ratio, molecular weight and the nature of terminal groups enables precise control over drug release profile [ 25 ]. In general, therapeutics can be orally released in a controlled manner via transport through the polymer or be triggered in response to environmental stimuli or released during polymer dissolution [ 26 ]. In addition, the bioavailability of hydrophobic drugs can be improved due to the hydrophilic property of PLGA, thus enhancing oral bioavailability [ 27 , 28 ].
Polymeric micro- and nanoparticles can be administered as

oral, injectable and inhalable drug delivery systems, and can be fabricated using various methods including solvent evaporation [ 29 ], emulsion [ 30 , 31 ], spray-drying [ 32 , 33 ], porous glass membrane emulsification [ 34 ] and coacervation [ 35 ]. However, most of these conventional approaches suffer from some limitations such as broad particle-size distribution, low drug encapsulation efficiency [ 36 ], particle agglomeration and difficulties for incorporation of hydrophilic drugs [ 30 , 37 ]. Moreover, non-degradable surfactants are employed in some of the approaches which might cause unwanted effects in pharmaceutical applications. In addition, for highly sensitive and active pharmaceutical ingredients, inactivation or degradation is possible owing to high temperature, organic solvent exposure and high shear stress during the particle fabrication process [ 31 , 38 ]. Electrohydrodynamic atomization (EHDA), also known as the electrospray technique, is believed to overcome the drawbacks associated with conventional techniques and has been deliberately chosen as a method for the fabrication of bioactive molecule loaded micro and nanostructured materials in the application of drug delivery [ [39] , [40] , [41] ]. In this method, drug molecules are firstly incorporated into the polymer solution and infused through a capillary nozzle. A high voltage is applied to atomize the solution and the electric charge generated on the droplet, competing with the surface tension and resulting in near-monodisperse microdroplets in the cone-jet mode and further solidify to form microspheres in a single-step process. EHDA also has the potential to avoid particle aggregation due to self-dispersing of the droplets owing to Coulombic repulsion [ 42 ]. Moreover, electrospraying has the advantage of high drug encapsulation efficiency [ 43 ] and can enable encapsulation of both hydrophobic and hydrophilic drugs [ 28 , 39 ].
In this work, electrospraying was utilized to produce progesterone loaded PLGA particles. Many studies have investigated the formulations of PLGA particles for controlled release of different pharmaceuticals [ 16 , 18 , 25 , 43 , 44 ] and encapsulated hydrophobic drugs within different polymer matrices [ 11 , 27 , 28 ]. Progesterone-loaded PLGA microparticles have been prepared by several researchers by use of different techniques, including microfluidic [ 26 ], hydrogel template [ 45 ] and

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Effects of Drug Physicochemical Properties on In-Situ Forming Implant Polymer Degradation and Drug Release Kinetics. — Jordan B Joiner et al., 2022

In-situ forming implants (ISFIs) are sustained drug delivery systems consisting of a biocompatible, water-miscible solvent such as N -methyl-2-pyrrolidone (NMP) or dimethyl sulfoxide (DMSO) and a biodegradable polymer, most commonly a polyester such as PLGA [ 1 , 2 , 3 , 4 ]. The water-miscible solvent is used to dissolve both the polymer and drug to form a liquid formulation that can be injected into the intramuscular or subcutaneous space [ 5 , 6 , 7 ]. The implant is formed via phase inversion, during which the water-miscible solvent diffuses from the ISFI into the aqueous injection site, leaving the precipitated polymer to trap the drug in a solid matrix [ 8 , 9 ]. Phase inversion kinetics have a strong effect on drug burst release (within the first 24 h) and can be tuned by changing system components, such as drug physicochemical properties or solvent, that alter the exchange between the aqueous environment and solvent [ 10 ]. The subsequent release rate is dictated by drug diffusion through the polymer matrix, and finally polymer degradation via hydrolysis [ 11 ]. Polymer hydrolysis rate is influenced by backbone structure, MW, crystallinity, hydrophilicity, microstructure, and material processing [ 12 , 13 ]. As a co-polymer, PLGA degradation can be fine-tuned by varying the ratio of lactic acid:glycolic acid, MW, molecular structure (linear vs. branched) and end group [ 14 , 15 ]. Previously, we showed that in vitro release of small-molecule drugs, MK-2048, and dolutegravir, from ISFI formulations can be tuned by altering the ratio of PLGA to NMP in the ISFI formulation [ 16 ]. The degradation byproducts of PLGA (lactic acid and glycolic acid) are non-toxic and known to be cleared through normal metabolic pathways, eliminating the need for surgical removal of the ISFI once all of the drug has been released [ 17 ].
ISFIs are an ideal drug delivery system for various applications, especially for those requiring frequent dosing regimens, because they are simple to manufacture, stable for months at room temperature, facile to administer, and can be removed to terminate treatment if required [ 16 ]. Although ISFIs

are a ternary system, it is crucial to consider the impact that each formulation component has on drug release kinetics to achieve optimal therapeutic effects. Formulation development typically requires performing multiple studies using a wide range of polymer–solvent systems until the desired drug release is achieved, which is time-consuming and costly [ 18 , 19 ]. Factors that affect in vitro drug release from ISFIs include drug loading, drug and polymer properties, polymer/solvent ratio, ISFI manufacturing process, type of release media used, and injection volume of ISFI [ 20 ]. Formulation work can be improved by rationally designing drug release studies based on known physicochemical properties of drugs. Two physicochemical properties, pKa, a measure of acidity, and logP, a measure of hydrophilicity, are particularly important because the degradation of PLGA has been shown to occur rapidly in acidic environments and because the drug’s affinity for the aqueous environment dictates whether it will interact more with the solvent or the polymer [ 21 , 22 ]. Here, we performed a systematic in vitro drug release study over 90 days using antiviral and chemotherapeutic drugs with a wide range of logP and pKa values with a constant formulation of 50 mg/g API in 1:2 w / w PLGA:NMP in order to study the effects of drug physicochemical properties on in vitro drug release kinetics and PLGA degradation. The future goal of this study is to build a predictive model for drug release from PLGA ISFIs at relevant biological conditions. PLGA was chosen as the polymer in these ISFI formulations because it is biodegradable, tunable, and has been previously used in Food and Drug Administration (FDA)-approved ISFI systems. The long-term benefits of this study are three-fold: (1) the development of a predictive drug-release model will greatly benefit researchers in academia and industry due to the rapid development of novel small-molecule drugs [ 23 ], (2) the development of a long-acting implant mitigates the need for daily dosing and improves drug compliance, and (3) the long shelf-life of ISFI formulations when stored at room temperature will allow for distribution of therapies to countries where access to cold chain storage and distribution are limited [ 16 ].
50:50 Poly(DL-lactide- co -glycolide) ester terminated (PLGA) was purchased from L

ACTEL (Birmingham, AL; Cat. No. B6010-1P, Lot# A17-142, weight average MW, 27.2 kDa, intrinsic viscosity (i.v.) 0.38, polydispersity index (PDI) 1.81). N -methyl-2-pyrrolidone (NMP, (USP)) was received from ASHLAND (Wilmington, DE, USA, Product Code 830697, 100%NMP). Dolutegravir base (DTG), rilpivirine base (RPV), darunavir base (DRV), ritonavir base (RTV), etravirine base (ETV), efavirenz base (EFV), raltegravir base (RAL) and idarubicin hydrochloride salt (IDA) were purchased from Selleckchem (Houston, TX, USA). Gemcitabine base (GEM) was purchased from LC Laboratories (Woburn, MA, USA). Lamivudine base (3TC) was purchased from Fisher Scientific (Pittsburgh, PA, USA). Zidovudine base (ZDV) was purchased from MedChemExpress (Monmouth Junction, NJ, USA). 5-fluorouracil base (5FU), phosphate-buffered saline (0.01 M PBS, pH 7.4), solutol-HS, HPLC grade acetonitrile (ACN), and water were purchased from Sigma Aldrich (St. Louis, MO, USA).
Reverse-phase HPLC analyses were developed and validated with a Thermo Finnigan Surveyor HPLC (Thermo Finnigan, San Jośe, CA, USA) equipped with a Photodiode Array (PDA) Plus Detector, LC pump plus, and autosampler [ 16 ]. Sample analyses were carried out on an Intersil, ODS-3 column (4 µm, 4.6 Å ~ 150 nm (GL Sciences, Torrance, CA, USA) stationary phase maintained at 40 °C, with a flow rate of 1.0 mL/min with a 25 µL sample injection. A mobile phase of H 2 O:ACN 95:5 v / v and 0.1% trifluoroacetic acid (TFA)

Mathematical modeling of drug delivery from autocatalytically degradable PLGA microspheres--a review — Ashlee N Ford Versypt et al., 2013

1.

Introduction

Introduction

Poly(lactic-co-glycolic acid) (PLGA) microspheres are controlledrelease drug delivery alternatives to conventional drug therapy regimens. By releasing drug molecules in a controlled manner over extended periods of time from a single administration, controlled-release systems have the potential to maintain drug concentrations within target ranges, diminish side effects caused by concentration extremes and repeated administrations, and improve patient compliance as compared to conventional regimens. PLGA microspheres have been extensively studied for controlled-release drug delivery mainly because of the biodegradable and bioabsorbable qualities that allow for the passive degradation of the polymer in aqueous environments such as living tissues and for the resorption of degradation products into the surrounding media . Despite these advantages, the implementation of controlled-release drug delivery devices composed of PLGA microspheres for human patients has been gradual; the characterization and design of the microspheres depends heavily on trial-and-error experiments, and the interplay between complex phenomena that contribute to the drug release is still being deciphered.
Several processes contribute to the overall kinetics of drug release from PLGA microspheres including chemical degradation of the polymer by autocatalytic ester hydrolysis, polymer erosion, evolution of pore structure as a result of mass erosion, and diffusive transport of the drug through the polymer matrix and the aqueous pore structure . In the present work, the term degradation refers to the process through which the polymer chain bonds are hydrolyzed to form oligomers and monomers. The term erosion refers to the loss of mass due to diffusion of water-soluble, small oligomers and monomers out of the polymer matrix. The definitions of degradation and erosion are the same as those given by Gopferich and have been widely adopted in the literature.
Three main phenomena-PLGA degradation, PLGA erosion, and drug transport-are discussed in Section 2, and mathematical models that mechanistically address these phenomena and the interactions between them are described in Section 3. The coupling between the three phenomena is important for understanding how one of the three may dominate or work in conjunction with the others under different conditions. The autocatalytic degradation mechanism may accelerate the degradation and erosion in the center of microspheres and enhance size-dependent drug transport. The complex effects of autocatalysis are difficult to predict without understanding of the

relative strengths of the phenomena and their dynamics.
Mathematical models can reduce the number of experiments needed to probe different conditions and designs and to deepen the understanding of the physical and chemical mechanisms of drug release, particularly when the effects of different phenomena are coupled. Empirical or correlative mathematical models, which are commonly applied in the drug delivery field, have very limited predictive capability outside of the specific experimental conditions used to fit parameters in the models . In contrast, mechanistic mathematical models aim to account for the physical and chemical phenomena that contribute to the overall drug release kinetics and are applicable over a wide range of conditions to be used in the model-based design of microspheres to produce desired release profiles (e.g., constant rate of release for uniform therapeutic dosage). Here, only mechanistic models and hybrid empirical and mechanistic models are addressed.

Background concepts

Polymeric drug delivery can be categorized based on the mechanisms of drug release : diffusion-controlled systems (diffusion from non-degrading polymers), swelling-controlled systems (enhanced diffusion from polymers that swell in aqueous media), and erosion-controlled systems (release as a result of degradation and erosion of polymers). For biodegradable polyesters such as PLGA, drug release occurs through a combination of degradation and erosion of polymer and transport of drug and is classified as being erosioncontrolled. In this section we overview the mechanisms of each of these processes for erosion-controlled drug release from PLGA microspheres and how their effects interact.

PLGA degradation

PLGA is a poly(a-hydroxy-ester) (see Fig. 1) that is depolymerized in the presence of water. The hydrolysis reaction cleaves the ester bonds of polymer chains. The reaction can be catalyzed by acids or bases, but experimental data on the acidic local pH within PLGA particles suggest that only the reaction mechanism catalyzed by acid is relevant. The acid-catalyzed ester hydrolysis proceeds by the bimolecular, acyl-oxygen cleavage A AC 2 mechanism [19,20] summarized by
n 1/4 2; 3; ... and m 1/4 1; 2; ...; n-1;

d1:1Th

where P n , P m , and P n - m denote polymer chains having degrees of polymerization n,

m, and n - m, respectively, and H + is the acid catalyst. The source of the acid catalyst can be external from strong acid in the medium (non-autocatalytic reaction) or internal from the carboxylic acid end groups of the polymer chains (autocatalytic reaction) . In autocatalysis, the reaction product catalyzes further hydrolysis in the manner
where A is water and B is acidic polymer chains in the context of PLGA degradation.

PLGA erosion

Polymer erosion is classified as surface-eroding or bulk-eroding [4,5,22,23]. For surface-eroding polymers such as polyanhydrides, the rate of polymer degradation at the surface is faster than the rate of penetration of water from bodily fluids in vivo or from the buffer medium in vitro into the polymer bulk. Surface-eroding polymers react from the surface inward. Bulk-eroding polymers exhibit a faster rate of water penetration than the rate of polymer degradation. The degradation and erosion in bulk-eroding polymers occurs throughout the polymer bulk. PLGA is a bulk-eroding polymer at the length scales used in drug delivery microspheres (10s to 100s of microns) as the hydration time scale is on the order of a few minutes compared to weeks or months for degradation .
Erosion depends on the degradation, dissolution, and diffusion processes . For PLGA, the dissolution of water-soluble oligomers up to nonamers and of drug molecules is often assumed to occur faster than diffusion and polymer degradation in many mathematical models and is neglected. A few models propose that dissolution is rate-limiting for PLGA oligomers .

Drug transport

An "initial burst" of drug release often occurs wherein a significant percentage of the drug is released during the early stage of the release process. This effect has been reported for many formulations of PLGA microspheres. The initial burst can be diminished or eliminated by adjusting the fabrication technique . The subsequent controlled release of drug molecules from PLGA microspheres depends on the transport properties of the drug and the dynamic conditions of the degrading polymer. The drug compound in PLGA microspheres may be released by some combination of diffusion through the polymer matrix, diffusion through aqueous pores, and dissolution coincident with polymer dissolution . Diffusion through the dense polymer matrix is possible but is

A Computational Model for Drug Release from PLGA Implant — Miljan Milosevic et al., 2018

1. IntroductionThe encapsulation and controllable release of drugs, as well as achieving enhanced therapeutic effects in drug delivery systems, were the subjects of investigation for a number of authors in the past (e.g., [1,2,3]). Among these drug-delivery systems, electrospun nanofiber mats are promising as drug carriers which offer site-specific delivery of drugs to the target in human body, and may be used for wound healing and cancer therapy [4,5,6,7,8]. Electrospinning is a technique that utilizes the electric force to drive the spinning process and to produce polymer fibers [9,10,11], and is capable of producing fibers with diameters in the nanometer range (10–1000 nm). Nanofibers obtained by electrospining are structurally homogeneous and are unlikely to encapsulate bioactive agents as nanoscaled particles. Recently, electrospinning of emulsions produced composite nanofibers with nanoscaled drug particles surrounded/coated by emulsifiers/surfactants and impregnated in biocompatible and/or biodegradable polymers [12]. Such types of composite nanofiber mats play the role of a controllable drug encapsulation/release vehicle.The most commonly used biodegradable synthetic polymers for three-dimensional (3D) scaffolds in tissue engineering are saturated poly(α-hydroxy esters), including poly(lactic acid) (PLA) and poly(glycolic acid) (PGA), as well as poly(lactic-co-glycolide) (PLGA) copolymers [13,14]. PLGA has been well recognized for its suitability in drug delivery due to its good biocompatibility and ability to achieve complete drug release as a result of degradation and erosion of the polymer matrix. PLGA is a linear copolymer that can be prepared at different ratios between its constituent monomers, lactic (LA) and glycolic acid (GA). Depending on the ratio of lactide to glycolide used for the polymerization, different forms of PLGA can be obtained. The degradation of PLGA copolymer is the collective process of bulk diffusion, surface diffusion, bulk erosion, and surface erosion. Since there are many variables that influence the degradation process, the release rate pattern is often unpredictable. The release of a drug from the homogeneously degrading

matrix is more complicated. Polymer composition is the most important factor for hydrophilicity and rate of degradation of a delivery matrix. Systematic studies of polymer composition with its degradation [15,16] showed that an increase in glycolic acid percentage accelerates the weight loss of polymer. It was shown that PLGA 50:50 (PLA/PGA) exhibited a faster degradation than PLGA 65:35 due to preferential degradation of glycolic acid proportion assigned by higher hydrophilicity. Subsequently, PLGA 65:35 shows faster degradation than PLGA 75:25, and so does PLGA 75:25 compared to PLGA 85:15 [17]. The absolute value of the degradation rate increases with the glycolic acid proportion. The amount of glycolic acid is a critical parameter in tuning the hydrophilicity of the matrix and, thus, the degradation and drug release rate [18].Modeling the PLGA degradation and erosion is a prerequisite for drug release modeling, and mechanistic approaches are most commonly employed [19,20]. Accompanied and facilitated by PLGA degradation and erosion, drug release has a significant impact by changing the properties of the polymer matrix (porosity and PLGA MW), and such factors need to be captured in the diffusion drug transport models. The hydrophobicity of electrospun nanofiber mats could play an important role in the overall performances as tissue engineering scaffolds. While macromolecular hydrophilic drugs are limited by diffusion through the pore space, relatively smaller hydrophobic drugs could diffuse through both the PLGA matrix and the pore space [21].It is a challenge to adequately model through numerical methods the process of drug release from fibers to the surrounding medium, with taking into account transport conditions within fibers (including degradation), in the medium, and at the interface between fibers and the surroundings. Using continuum elements for modeling fibers would require a huge effort for the finite element (FE) model generation and lead to an enormous number of equations, therefore preventing implementation to practical problems. In order to have a robust model, feasible for practical use, we have introduced two approaches here: (1) The use of a radial 1D finite element which replaces a detailed modeling of fibers by continuum elements [22], and (2) a model with fibers represented by a continuum according to the smeared concept introduced in References [23,24,25]. The second model is particularly attractive, since

it does not need any 1D finite element meshing for fiber representation.In the next section, Materials and Methods, electrospining methodology and the drug release of RhB from PLGA1 (65:35) and PLGA2 (50:50) nanofiber mats are investigated. Next, we present fundamental equations of the radial 1D element and the equation of degradation implemented into our model, including the hydrophobic effects. This is followed by the formulation of the smeared model for the fiber network. Finally, we demonstrate the applicability and accuracy of the computational models by presenting both numerical and experimental results. 2. Materials and MethodsHere are described the materials used for electrospining, as well as the electrospining procedure and the design of simple experiments for drug release. 2.1. MaterialsPoly(d,l-lactide-co-glycolide) PLGA1 (average molecular weight MW 40,000–75,000 g/mol) with the mass ratio of lactide:glycolide units being (65:35), poly(d,l-lactide-co-glycolide) PLGA2 (average molecular weight MW 30,000–60,000 g/mol), with the mass ratio of lactide:glycolide units being (50:50), Rhodamine B (RhB), span-80, N,N-dimethylformamide (DMF) and chloroform (CHCl3), were purchased from Sigma-Aldrich Co. (Milwaukee, WI, USA). The chemicals were used without further purification. Phosphate buffered saline solution (PBS) was made by dissolving one tablet of PBS, supplied by Fisher Scientific (Hampton, NH, USA) in 200 mL of distilled water. 2.1.1. Preparation of PLGA Nanofibers Produced via Emulsion ElectrospinningTo prepare the emulsion containing 24 wt.% PLGA and 0.1 wt.% Rhodamine B, PLGA1 and PLGA2 (3.0 g) were initially dissolved in the mixture of solvents chloroform/DMF (8.25/2.75 g) and magnetically stirred at 200 rpm at room temperature for 24 h. After that, span-80 (50.0 mg) was added to this polymer solutions, followed by the addition of 5 wt.%