Microfluidic assisted synthesis of PLGA drug delivery systems

Poly(lactic-co-glycolic acid) (PLGA) is a biocompatible and biodegradable polymer that recently attracted attention for use as part of drug delivery systems (DDS). In this context, there is an emerging need for a rapid, reliable and reproducible method of synthesis. Here, microfluidic systems provide great opportunities for synthesizing carriers in a tightly controlled manner and with low consumption of materials, energy and time. These miniature devices have been the focus of recent research since they can address the challenges inherent to the bulk system, e.g. low drug loading efficiency and encapsulation, broad size distribution and burst initial release. In this article, we provide an overview of current microfluidic systems used in drug delivery production, with a special focus on PLGA-based DDS. In this context, we highlight the advantages associated with the use of microchip systems in the fabrication of nanoparticles (NPs) and microparticles (MPs), e.g. in achieving complex morphologies. Furthermore, we discuss the challenges for selecting proper microfluidics for targeted DDS production in a translational setting and introduce strategies that are used to overcome microfluidics shortcomings, like low throughput for production.


Introduction
Drug delivery systems (DDS) aim to administer an optimum dosage of drug to the body in order to treat the disease or provide relief pain. 1 DDS are specially designed to deliver drugs to desired sites in a sustainable manner. 2 Recently, growing research has been started to nd systems that have the features as well as the facility and reproducibility for fabrication methods. 3 Drug delivery systems include micro-and nanomaterials that have applications in tissue engineering, therapy, and even in diagnosis and imaging. For this reason therapeutics with modied activities and features have been the focal points of this extensive research. Polymeric carriers with the capability to tailor and engineer structures are among the most appealing materials for researchers. 4 Delivery of hydrophilic and hydrophobic drugs is achievable via designed polymeric DDS. Problems related to solubility, degradation, and toxicity of various polymers have led researchers towards biocompatible and biodegradable polymers such as aliphatic polyesters. 4,5 PLGA is one of the best characterized biodegradable and biocompatible copolymers that decomposes to nontoxic products (H 2 O and CO 2 ) that are eliminated from the body through the Krebs cycle. Typically, PLGA is produced by a catalyzed ringopening copolymerization of lactic acid (LA) and glycolic acid (GA). Poly(glycolic acid) (PGA) is a crystalline, hydrophilic polymer with low water solubility and fast degradation rate under physiological conditions. On the contrary, poly(lactic acid) (PLA) is a stiff, hydrophobic polymer with low mechanical strength. As a copolymer, PLGA inherits the intrinsic properties of its constitutional monomers. PLGA properties can be tailored for specic applications just by varying the ratio between LA and GA monomers. Due to this phenomenon, PLGA represents a good material for drug delivery systems of many therapeutic agents (e.g. chemotherapeutics, as well as antiseptic, anti-inammatory and antioxidant drugs and proteins). Some of these PLGA based DDS have been approved by the US Food and Drug Administration (FDA) or are in clinical phase trials (Fig. 1).
Drug-loaded PLGA microparticles (MPs) [6][7][8] and nanoparticles (NPs) [9][10][11][12] can be synthesized via various bulk methods such as salting out, membrane emulsication, single/double-emulsion, nanoprecipitation. Particles produced by conventional bulk methods usually suffer from high batch-to-batch variation and polydispersity. These drawbacks arise due to uncontrollable synthesis method. 13 For example, all stages of NPs formation including nucleation, growth, and agglomeration take place simultaneously which leads to polydisperse particle formation. 14 Production process needs good control over surface charge, size and size distribution since these parameters control the drug release rate. In microuidics the mixing rate, heat, and mass transfer are more precise, synthesis in these miniature devices is more controlled. Moreover, shorter mixing and process time provide another advantage of these small devices; as lower material consumption. These aspects of microuidics are very advantageous comparing with conventional drug production methods. Therefore these miniature devices can be used to address the conventional methods issues.
The aim of this review article is to collectively encompass the PLGA DDS produced in microuidics, to address the impact of solvent and microuidic system in the size and properties of PLGA-based drug delivery systems and pave the way for researchers to choose better system to achieve their goal in the production of drug delivery systems with sophisticated and precise properties.

Microfluidics
Microuidics have been dened as devices in which small volumes (micro-or nanoliter) of liquids are processed or manipulated in microchannels to achieve better control on mixing, heat and/or mass transport. 17 The microchannels are made of various materials such as polymers (e.g. polydimethylsiloxane (PDMS) or polyimide), metal (aluminum) and glass capillaries. As microuidics synthesis allows to tightly control the properties of particles, the technology offers a broad range of advantages over the conventional bulk methods (Table  1). [18][19][20][21] Microuidics methods can be divided into two main categories based on the ow conguration in the microchannels: droplet-based (segmented) and continues microuidics.

Droplet-based microuidics
Droplet-based microuidics are used to synthesize microdroplets, emulsions, and microparticles as well as  Table 1 The advantages and disadvantages of microfluidics for synthesis 3,13,22,23 Advantages Disadvantages microreactors for nanomaterial production. In the current section, we will provide a brief description of the rules of thumb and principles of droplet generation. Droplets form because of the instability in the inner ow which breaks into drops. Many parameters are important in the droplet formation, but crucial ones are channel geometry, ow rate, uid viscosity and surfactant addition. For example, channel design (e.g. contraction or the method uid ows come into contact) and channel diameter are determinant in the droplet formation phenomena and properties. Furthermore, uid properties such as viscosities and the presence of surfactants are effective parameters in the viscous shear forces, which break the inner stream into droplets.
Droplets can form in various regimes of ow. 24 Three main regimes are dripping, jetting and squeezing. The dripping mode has been observed in low ow rates and by an increase in the ow rate, it changes to jetting mode. 25 Dripping regime produces droplets with narrow size distribution (Fig. 2I) while jetting mode produces polydisperse droplets. 26 In the jetting regime, droplets are small with higher surface-to-volume ratio 27 and form far from channels' exit ( Fig. 2II). 25 In the squeezing mode (Fig. 2III), droplets start to grow and plug the continuous phase and consequently by the increase in the pressure of continuous phase they breaks off. Therefore, the squeezing mode is characterized by a uctuation in the pressure of uids.
Droplet generation as well as mixing in microuidics can be performed by active and passive methods. 24 In the active method, an external force (e.g. magnetic, electric and etc.) is applied to facilitate the droplet formation. In contrast, in the passive mode, two or more immiscible uids come to contact in a junction and droplets form depending on the properties of the uid (ow rate ratios, ow conditions, and the geometry of the device). 25 Various geometries have been evaluated to promote droplet generation with single or multiple cores. 27 Based on the ow contact, geometries are classied into three main categories: co-ow, cross-ow, and ow-focusing ( Fig. 3I-III). 28 In cross-ow; continuous and dispersed phases meet in a junction with various angles (Fig. 3I) and both stream properties affect the ow and droplets form. Based on the angle of contact, there are various T- 90 < a, Q < 180 ). Additionally, designs with more than two inlets have been used as double T-, V-and K-junction. 25 Out of all junctions, the T-junction is most frequently used since it produces droplets with narrow size distribution. 25,27 In a co-ow geometry, dispersed phase ows from an inner channel coaxially with the continuous phase in the outer channel with higher ow rate (Fig. 3II). Finally, in the ow-focusing method (Fig. 3III) two phases (dispersed and continuous) ow coaxially and pass through a region with contraction. Regarding the interfacial properties of the phases, at this point, the phase with lower ow rate breaks up into droplets and emulsion forms. Moreover, considering the number of owing phases, it can produce single or double emulsions.
Aforementioned designs produce single emulsions. Double emulsion can be produced with a combination of them, such as T-junctions in cascade in glass capillary (Fig. 3IV), three co-axial channels with one contraction region (Fig. 3V), multiple owfocusing designs in series (Fig. 3VI), and combination of a Tjunction and a ow-focus (Fig. 3VII). Many researchers have tried to produce multi emulsions in one step (Fig. 3VII), 26,29 but it is difficult to control ow and the conguration becomes very complex. As it is seen in Fig. 3, microuidics makes it possible and feasible to generate droplets with varying core size ( Fig. 3IX and X) and one core with multiple shells (Fig. 3IX and X). Such droplet morphologies are the basis of microcapsules, core-shell MPs, polymersomes, lipid vesicles, etc. 2

Continuous phase ow microuidics
In this type, two or more uids ow side-by-side in microchannels without segmentation or breakup. Researchers try to adopt continuous phase ow for material synthesis due to its reduced mixing time. 30 A reduced mixing time, resulting from the compression by the outer uid, is very important in many NPs syntheses since it provides a homogenous condition for NPs formation. As a result, NPs have a more narrow size distribution. Furthermore, uniform concentration, heat, and uid prole take place in inner uid and away from channel walls which prevent particle generation close to the channel wall and as a result, it reduces channel clogging. 31 There are two types of devices for continuous phase ow microuidics: the coaxial tube devices which are widely used for inorganic synthesis 32-34 and the hydrodynamic focusing (HF) devices. The latter are very exible in design and various subtypes have been recorded (Fig. 4II-IV) based on the number of uids and angles of contact in the focus point. Fig. 4II shows the simplest HF design with one stream compressed between two streams in various contact angles (a # 90 ). Multiple HF is possible in a sequential manner (III) or one contraction point (IV). Depending on two angles (a & b) and distances (d & d 0 ) this conguration can be used to improve ow stability or avoid central synthesis regime from channels wall. 31 Recently, 3D designs of HF devices have been investigated in nanomaterial synthesis, in which the inner uid squeezes between an outer ow horizontally and vertically. [35][36][37] Coaxial tube designs ( Fig. 4I) were also considered as 3D HF with circular cross section. 27,38 3D HF in microchannels with the rectangular crosssection is more complicated and difficult to achieve stable ow. Since in these 3D HF designs particle formation takes place away from channel walls, it prevents clogging and also aggregation of particles.

Microuidics systems used in PLGA drug delivery systems
In the last two decades, PLGA-based drug delivery systems are being produced in microuidics. There are several reports of PLGA-based MPs, NPs and microbers produced using this technology. A wide range of drugs has been loaded into PLGAbased MPs and NPs (e.g., bupivacaine, 39 risperidone, 40 ibuprofen, 41 paclitaxel (PTX), 42 doxorubicin (DOX), 43 camptothecin 43 and etc.) in various microchannels such as PDMS microuidics, 39 glass capillary, 43 phenol formaldehyde resinbased microuidic chip, 44 aluminum 22 and silicon. 45 However, production of microbers in the microuidics is not the commonly used synthesis method, especially for drug delivery purposes. The method has been employed to produce micro-bers for tissue engineering scaffolds. 46,47 However, microbers produced by this method have shortcomings to be used in DDS such as the presence of voids in the microbers structure 48,49 and hydrogel nature of the produced bers due to solidication process. 50 Considering the microuidics conguration, various methods of ow and mixing have been employed for NPs production in micro-channels such as segmented or continuous phase ow. Inorganic nanomaterials are synthesized in segmented ow or droplets based microuidics. [51][52][53][54][55][56][57] In such systems, micro-scale droplets serve as micro-reactors. In the case of PLGA, use of droplet-based microuidics results into micron size particles. MPs have been synthesized via reactions or phenomena that turn emulsions and droplet templates into particles. These phenomena include solvent evaporation or extraction which is not easily achievable in nano-scale.
Furthermore, according to the reports, droplet sizes are proportional to the size of the channels. Nano-scale droplets are necessary for NPs production. On the other hand, manipulation and pumping of uids in such small size channels and capillaries need a large amount of power. Additionally, it is almost impossible to online characterize and control the uids and droplets properties in nano-scale.
Literature survey led us to the idea represented in Fig. 5, i.e. the use of droplet-based microuidics produces PLGA MPs and particle synthesis in the continuous microuidics will end up at nano-scale. In Fig. 5 each line with a number over it indicates the size range reported by the associated reference article. It can be seen that the type of microuidics is very important in the scale of the drug delivery system. In more detail, except a few papers, there are not many reports on the PLGA nanoparticles production via droplet microuidics. 26 For instance, Lee et al. succeeded in PLGA micro/nanosphere production through droplet-based microuidics. They investigated solvent evaporation and extraction effect in the nal particles sizes. Fast evaporation of the PLGA droplets in dimethyl carbonate (DMC) produced PLGA MPs (3 to 30 mm). 58 However, they generated PLGA droplets in dimethyl sulfoxide (DMSO) in silicon oil continuous phase and aer infusion with water droplets, solvent extraction led to nano-size (70 to 500 nm) PLGA particle formation. In both methods, particle size increases with the higher concentration of PLGA. In another report, Janus PLGA NPs have been synthesized with T-junction uidic device consisted of two stainless steel capillaries that enter in a transparent plastic tube. From one inlet PLGA/PLA and Nile red mixture in dimethylformamide (DMF) and the other PLGA and rhodamine 6G in acetone added (with 100 mL h À1 ow rate) to an aqueous solution containing 1% PVA (ow rate 10 mL min À1 ). The authors used the same system to load PTX and DOX in PLGA NPs and claimed that the approach overcomes both drugs limitations and offers a high yield of dual-loaded NPs. 59 Fig. 5 reveals that directly the microuidic type and indirectly, solvent miscibility, ow rates, and their ratio, the synthesis processes are very crucial in the nal characteristics of the drug delivery system and efficiency. It appears that the particle formation process is critical in the nal size and size distribution of the carrier. For example, PLGA-PEG solution in a water miscible solvent (acetonitrile, DMSO, DMF) focused with aqueous solution gives rise to self-assembly of polymer chains into polymeric NPs and micelles. 14,60,61 However, the similar polymeric composition in a water immiscible or semimiscible solvent (DCM) produces droplets which need a further step to remove the solvent and produce MPs. 62 For example, change in the solvent from DCM to DMSO and also infusion with water droplets resulted in the sub-micron particles. Although they claimed the droplet-based microuidic produced nano-scale PLGA particles, the infusion of polymer mixture (water-miscible solvent) with water droplets produced NPs which the precipitation takes place within the droplets. 58 Additionally, NPs fabricated using DMF as an organic solvent in the T-junction droplet producing a uidic chip with very low ow ratio. 59 It can be concluded that precursor's properties and the conguration of streams in microchannels have the decisive inuence in the formation process. Although the device resembles droplet-based microuidics, comparison of dispersed and continues ow rates brings dropwise nanoprecipitation (in the bulk method) into mind. Moreover, for NPs production solvent can signicantly affect the properties of drug loaded systems (size and size distribution). 63,64 Solvents with higher diffusion coefficient produce NPs with smaller size and narrower size distribution. 64 PLGA NPs produced in continuous microuidic are in submicron size with low encapsulation efficiency and drug loading. This challenge is associated with water miscibility of the solvents. During the NPs formation, as a result of solvent displacement, polymer chain come together and form smaller particles. However, a huge part of agents is lost at the same time during solvent displacement. Recently, in an interesting approach, Xu et al. used glass capillary droplet-based micro-uidics in which emulsion was generated by the combination of two solvents; DMSO and DCM. 65 Although the organic phase forms droplets, one of the solvents (DMSO) displaces into aqueous phase while DCM entraps the drug (DOX) within the droplets during NPs formation, avoiding drug lost. In agreement with the statement, encapsulation efficiency (48.5, 49.9, and 56.9%) and drug loading (9.7, 10.0, and 11.4%) increase in NPs with increase in the DCM ratio in solvent mixture (V DCM / V DMSO : 0, 0.05, 0.1). Taken together, the particle formation process is the basic determining factor of the DDS size. In the continuous micro-uidics, PLGA particles form by a nanoprecipitation process in the interface of water-miscible organic solution (middle stream) with the aqueous stream from both side and even sometimes from up and downside. This process is very fast and happens in nano-scale and eventually NPs form. However, in the case of droplet microuidics, PLGA polymer with drug or agent dissolved in a water-immiscible (or partially miscible) solvent produce template emulsion or droplets that typically evaporated to remove the solvent and produce MPs. As a result, the organic phase solvent and microuidics conguration regulate the particle formation process, pace and consequently the produced particle size.

PLGA-based nanoparticles
As it has been mentioned in the previous section, PLGA NPs synthesis in continuous ow microuidics is accomplished by the nanoprecipitation process between two phases that ow alongside. In this process, a material solution containing polymer and drugs in a water-miscible solvent (e.g. acetone, acetonitrile, ethanol, or methanol) is compressed within a nonsolvent phase such as an aqueous solution containing surfactants. 66 The solvent is miscible in non-solvent and transfers between two phases which leads to NPs formation. 30 Various kind of microuidics (2D HF to 3D, laminar ow to turbulent jet) has been used to produce PLGA-based NPs.
The simple yet convenient 2D HF type allows fabrication of multi-drug loaded PLGA NPs with desired properties. For example, bisphosphonate conjugated PLGA (BP-PLGA) NPs loaded with superparamagnetic iron oxide nanoparticles (SPIONs) and PTX are produced in a PDMS based chip (Fig. 6I) to be used in chemotherapy, hyperthermia, and MRI diagnosis. 67 Results show that NPs produced in microuidics (in the range of 40 to 100 nm) is smaller in comparison with bulk method ($120 nm). Furthermore, it is possible to control the properties of NPs with ow conditions in 2D HF microuidics. For example, by an increase in the ow rate ratio (the ratio of the organic phase containing NPs precursors to aqueous phase), nal NPs size increases (Fig. 6II). Additionally, this ratio can affect drug release prole, as in lower ratios, NPs are smaller and more compact, consequently, drug release is slower. Results from in vivo analysis with a bone metastasis model (MDA-MB-231) mice showed that tumor growth suppression and apoptosis level enhanced with targeted microuidic NPs. In a recent study, 68 it was shown that an increase in the ow ratio from 0.025 to 0.125 means size of non-targeted curcuminloaded PLGA NPs increase from about 30 nm up to 70 nm. The authors stated that the higher the ratio, the broader the size distribution of produced NPs. Results from in vitro analysis exhibited that compared with free curcumin, microuidically produced NPs enhance the antitumor activity of the drug toward leukemia Jurkat cells. As it's seen, ow ratio is an effective parameter in the synthesis process and lower ow ratio produces smaller NPs with narrow size distribution. However, lower ow ratio means the slower ow of NPs precursors and an undesirable consequence of slow ow is low efficiency which is the intrinsic characteristic of low ow rates.
Another effective parameter is mixing time and determined by the design of the microuidic chip. Mixing time controls the particle formation and also nal properties and the amount of NPs. In 2D HF microuidics with microchannels, ows are laminar; mixing is based on diffusion which takes place in the interface of phases. In order to investigate the effect of design on mixing time and also NPs properties, PLGA-DOX solution in the mixture of DMF and triuoroethanol (TFE) was introduced into three PDMS based chips: 2D at HF, 3D arc, and 3D double spiral. 69 Simulation results represented that the mixing time as a function of ow rate can be prolonged by an increase in ow rate. Moreover, 3D designs have a shorter mixing time due to the shortened mixing distance. For instance, in the same ow rate (2.5 mL h À1 ) mixing time decrease for 2D at HF, 3D arc and 3D double spiral (29, 16, and 14.5 ms, respectively). However, encapsulation efficiency (ca. 50%) and cellular uptake with MCF-7 and HeLa cells are reported only for 100 nm NPs produced in origami chip. DOXloaded PLGA NPs were more taken up with cancer cells and showed higher cytotoxicity compared with free DOX. In another conducted research, Liu et al. used 3D coaxial ow in the capillary glass to minimize the mixing time. They increased the ow rates which caused ow regime transfer from laminar to turbulent jet. 63 An additional distinguishing feature of their design is that organic phase (PLGA solution and PTX) ows as outer uid near the wall of the channel and aqueous phase ows in the central glass capillary. Higher drug loading and encapsulation efficiency (Fig. 6III) reported for microuidic NPs compared with the bulk method. The results attributed to the fact that volume ratio between polymer precursor and an aqueous solution for the microuidics system is xed and higher than bulk synthesis (the ratio is low and increases gradually during the process). PLGA NPs with the size ranges of 100-210 nm can be produced with mass production rates up to 242.8 g per day. In a similar conguration with one more capillary glass (Fig. 6IV) PTX and sorafenib (SFN) (anti-angiogenic drug) are assessed to be loaded in core-shell NPs. 70 Results declare that drug loading increases for PTX (from 6.7% to 42.6%) and SFN (from 6.2% to 45.2%) with a sequential conguration in comparison with single step process. As it's seen in Fig. 6V, the higher the ow rates (higher Reynolds number) the smaller the NPs size. Microchannel size is another effective parameter in the formation and properties of NPs. In larger diameter channels the ratio drops and NPs size increases. The smaller size of the NPs produced in microuidics is a result of a higher surface-tovolume ratio offered by these devices. In a 3D microuidics fabricated out of commercially stainless steel capillary with three inlets, dexamethasone and ribavirin encapsulated in NPs with size range 35-350 and 50-200, respectively. 71,72 According to the outcomes, by a decrease in the internal diameter of the channel from 600 to 130 mm, NPs size dropped from about 133 to 28 nm. Researchers mentioned that easily assembled device has the capability to be used in a series of parallel designs for mass production up to 2.4 kg NPs per day. 71

PEG-PLGA NPs
PEGylation is the process of polyethylene glycol (PEG) chains conjugation on the molecules or microstructures. The NPs consists of PLGA core and PEG chains on the surface offer a wide range of advantages; e.g. long circulation time due to immune system evade, small size around 20-100 nm, higher solubility, and stability, capability for drug encapsulation, good degradability, and biocompatibility. 73 PEG chain conjugation to PLGA has been investigated for a long time. PLGA-PEG copolymers assemble into NPs or micelles in the aqueous phase. Therefore, the precipitation in the aqueous phase to produce NPs is being used for a long time in the bulk method and it has been adapted to the microuidic system in the last decade.
For the rst time, Karnik et al. reported nanoprecipitation of PLGA 15k -PEG 4k in 2D HF microuidics. 14 They reported that NPs size and drug encapsulation efficiency are affected by ow rates and composition of phases, i.e. incorporation of PLGA increases encapsulation efficiency from 28% to 51%. Incorporation of PLGA or PLA increases the NPs size in bulk method. However, to prevent the size enlargement, PLA modied with prodrug (platinum(IV) [Pt(IV)]) has been incorporated to PLGA-PEG for drug release control. 74 The polymer solution used to load hydrophilic cisplatin and hydrophobic docetaxel into NPs in a 2D HF microuidic. Nanoprecipitation with identical condition produced smaller NPs in a microuidics ($100 nm) compared with the bulk method (greater than 150 nm). In a similar study, cisplatin prodrug (as conjugated to PLA backbone) and free irinotecan loaded into PLGA-PEG NPs to target prostate-specic membrane antigen (PSMA) overexpressing prostate cancer cell using S,S-2-(3-[5-amino-1-carboxypentyl]ureido)-pentanedioic acid ligand. 60 Addition of cisplatin-PLA to the solution of PLGA-PEG with irinotecan increased its encapsulation efficiency from 10% up to 44%. For both studies, results from in vitro analysis with LNCaP cells showed that the combination of drugs increases the cytotoxicity toward the cells compared with single drug-loaded NPs. In the following study, the authors used two-stage microuidics for mixing and production of doxorubicin-loaded PLGA-PEG targeted NPs in a fully integrated microuidic device. 61 They could produce a library of NPs with various surface properties, ligand densities, size, and molecular weight to evaluate in vitro and in vivo. Addition of 14% mole of targeted polymer to PLGA-PEG mixture increases the LNCaP cells uptake and tumor accumulate up to 3.5-fold compared with bare PLGA-PEG NPs. In order to promote the cellular uptake, pH-sensitive NPs have been produced in a similar 2D HF microuidics (Fig. 6VI) which can escape endo/lysosomes and overcome drug resistance. 75 The core-shell NPs consisted of DOX-loaded PLGA core and poly(ethylene glycol)-poly(2-(diisopropylamino)ethyl methacrylate) (PEG-b-PDPA) diblock copolymer shell. The authors claimed that upon the translocation in the acidic endocytic, the residual parts of shell produce positive charges over the PLGA core and help lysosomal escape. Results for ex vivo and in vivo analysis with MCF-7/ADR tumor-bearing mice reveal that in comparison with free DOX, the NPs signicantly suppress the drug-resistant tumor growth (Fig. 6VII and VIII).
To produce as small as possible NPs with homogeneous size, PLGA-PEG self-assembly investigated in 3D hydrodynamic focusing microuidics with sequential inlets. 35,36 Results disclosed that not only the concentration of polymer is important but also polymer molecular weight is a contributing factor in the nal size of NPs. For instance, with an increase in PLGA molecular weight (10 to 90 kDa) the produced NPs size increases ($26 to 150 nm). Moreover, in the same molecular weight (10 kDa) with increase in the concentration (10 to 50 mg mL À1 ) the NPs size increases (13 to 26 nm, respectively). The 3D micro-uidics has reduced the fouling and clogging since reaction takes place far away channels wall. 35 However, the 3D micro-uidics is complicated and it is difficult to achieve stable ow and reproducible manner. To increase the throughput, 8 parallel 3D devices in one chip, they reduced the batch time (for 25 mg) from 5 h to less than 20 minutes. 36 Low throughput of PEGylated PLGA NPs produced by these miniature devices is an important challenge and it has been tried to overcome in various research. In one attempt a 3 layer PDMS microuidics with 100 channels fabricated that produced methoxyl PEG-PLGA (MPEG-PLGA) up to 0.5-2.0 mL h À1 polymer ow rate with narrow size distribution. 76 Another 3D HF microuidics fabricated with parallel polyimide lms which can tolerate up to 16 MPa with high throughput up to 331 g per day of PEG-PLGA NPs. 77 Lim et al. 78 designed a turbulent jet micromixer with a higher ow rate and consequently high production more than 3 kg per day which is the highest throughput achieved up until now. 3

Lipid-PLGA NPs
Lipids are hydrophobic or amphiphilic molecules which also can be used to modify molecules. They have attracted attention for PLGA surface modication. 79 These core/shell structures have hydrophobic cores and hydrophilic tails of lipids. These structures are capable of hydrophobic drug loading and have prolonged circulation time compared to PLGA NPs. 79,80 Recently, lipid-PLGA NPs have been produced in the microuidic device to control the reaction and properties of the particles. In agreement with results, lipid-PLGA NPs with a smaller size ($62.5 and $87 nm) are produced in lower total ow rate (41, 246 mL h À1 , respectively). Moreover, cellular uptake evaluation with A375 cells (human melanoma cell line) indicated that smaller lipid-PLGA NPs are internalized more efficiently compared with larger counterparts. 81 Zhang et al. synthesized dual drug (DOX and combretastatin A4 (CA4)) loaded PLGA NPs with mono-and bi-layer lipid shells in a twostage HF microuidics. 82 Cellular uptake analysis with HeLa (cervical cancer cells) and HUVEC cells (Human umbilical vein endothelial cells) showed that NPs with monolayer lipid was taken up more than bilayer counterparts and even free drugs. Furthermore, similar ndings observed with in vivo and ex vivo analysis, i.e. monolayer NPs exhibited improved anticancer activity and faster tumor accumulation. The same procedure adapted to produce lipid-PLGA and lipid-water-PLGA NPs loaded with doxorubicin in PLGA core and CA4 in the shell, respectively. Rigidity analysis showed that NPs with a layer of water between polymer and lipid layers are more exible than the bi-layer counterparts. In vitro analysis (with HeLa and HUVEC cells) veried the results from a molecular dynamics simulation that revealed rigid NPs have enhanced cellular uptake compared with free drugs and exible NPs. 83 According to the ndings that rigid lipid-PLGA NPs exhibit higher cellular uptake, similar morphology adapted to load hydrophilic agent (siRNA) in water core, hydrophobic drug (DOX) in PLGA layer and lipid shell. The core/shell morphology enabled co-delivery of siMDR1 (the siRNA sequence against the multi-drug resistant protein) and doxorubicin and performed an enhanced gene knockdown efficiency compared with lipofectamine 2000 . In vitro analysis (Fig. 6IX) by MCF-7/ADR cells showed that cellular uptake of the lipid-PLGA NPs ($100%) is much higher than free DOX (30%). The NPs was surprisingly effective in tumor growth inhibition for mice treatment in comparison with and free drug, free gene. 84 Lipid shells not only used to load drugs also improved the stability of polymeric NPs but also used to enhance quantum dot (QD) nanocrystals hydrophilicity and biocompatibility in biological environments. In a conducted study, PLGA solution in acetonitrile focused in a 2D HF microuidics with lipid solution (aqueous solution of lecithin and 1,2-distearoyl-snglycero-3-phosphoethanolamine-N-[carboxy(polyethylene glycol)] (DSPE-PEG)) and aerward mixed in a Tesla microstructure to produce lipid-PLGA NPs. Additionally, in order to load quantum dots for diagnosis application, they used the same arrangement and solutions in which the aqueous phase contained lipophilic quantum dots (dissolved in tetrahydrofuran). 85 Results exhibited that rapid mixing in Tesla micromixers produces monodisperse NPs (35-180 nm) since it improves mixing efficiency. Results with various experiment revealed that lipid-PLGA NPs with 40 nm size have the most stable form and also lipid : PLGA (e.g. 1 : 10 to 1 : 1000) ratio is not effective in the size of NPs. Another imaging agent (gold nanocrystals (AuNCs)) with two therapeutic loaded in lipid-PLGA NPs. With an interesting approach, PLGA which was functionalized with AuNCs forms a hydrophobic core loaded with DOX and a lipid layer contains SRF, lipid shell composed of ordinary phospholipids and PEGylated phospholipids. 86 Another interesting point about the study is the use of a 3D microuidic chip with three inlets to increase the production and control the size of NPs. Drug release analysis reveals a sequential release of drugs; SRF release followed by DOX release and in vivo evaluations represent higher accumulation in the tumor site. Kim et al. 87 used the same chip to investigate the ow pattern and condition on the size and mass production of lipidic PLGA NPs. In the study, organic solutions (acetonitrile containing polymer) in the middle inlet and aqueous phase (lecithin and DSPE-PEG) in the outer inlets generate micro vortices. NPs sizes are affected by ow rates and Reynolds number, as in higher Reynolds numbers (Re ¼ 150, 75), NPs size decreases (55 nm and 81 nm, respectively). The authors claimed that their microuidics has higher productivity (up to 3 g h À1 ) which is 1000-fold of the conventional 2D HF microuidics. Gdowski et al. 88 used a micromixer with herringbone pattern (Fig. 6X) to promote the mixing of PLGA and curcumin in acetonitrile with an aqueous phase containing DSPE-PEG. They optimized NPs size to 102.11 nm with 4.4% drug loading and 58.8% encapsulation efficiency. However, in vivo assessment with mice bearing prostate cancer cells (C4-2B) showed that NPs are completely removed from the body aer 24 h (Fig. 6XI).

PLGA-based microparticles
Microparticles have great importance in biomedical applications because of their capability in the delivery of a broad range of drugs, higher encapsulation efficiency, controlled and stimuli release. 25,39,89,90 Due to the good properties of PLGA for biomedical applications, researchers have considered different geometries, chips, and congurations to produce PLGA-based microcarriers with tunable size and morphology. Table 2 summarizes PLGA MPs production in various microuidic systems with different agents to be used in drug delivery systems. The current section covers PLGA MPs produced in microuidic systems.
Typically, MPs in microuidics are produced via template droplets that turn into MPs through various reactions or in the case of PLGA, by solvent evaporation or diffusion. 25 Rapid processing in microuidics needs a volatile solvent to evaporate rapidly as well as dissolve organic component. For this aim, organic solvents such as DMC, chloroform, and toluene are frequently used to produce droplets in an aqueous continuous phase. The solvent is important since it can affect the production process or even the properties of the nal particles. For instance, encapsulation efficiency, an initial burst release of enoxacin (ENX)-loaded PLGA MPs synthesized in a PDMS microuidics is controlled by changing the solvent from DMC to dichloromethane (DCM). 100 MPs produced by PLGA dissolved in DMC exhibited higher encapsulation efficiency (56.5%) and initial burst release (14.8%) compared with DCM (encapsulation efficiency 15.4%, initial burst release: 12%). The concentration of the organic phase is effective in the size of particles. With the increase in PLGA concentration (1, 3, and 5 wt%) in DMC, MPs size increases from 140 to 160 mm.
Stability investigation of PLGA MPs has been carried out by various researchers. 101 Tu and Lee investigated the effect of PLGA composition and the pH of the inner phase of water-inoil-in-water (W/O/W) double emulsions. They showed that basic phase in double emulsion PLGA microcapsules enhances the surface activity and consequently improve colloidal stability. 102 Furthermore, it has been proven that monodisperse MPs produced in microuidics have better colloidal stability in aqueous dispersion when compared MPs produced a bulk method that tends to aggregate in storage. 23 Droplets and emulsions stability is also important in the nal characteristics of MPs which can be affected by the chip properties and conguration. For instance, PDMS is hydrophobic and it's difficult to produce droplets with varying hydrophobicity. Some modications have been proposed and used 103,104 such as immersion in the poly(vinyl alcohol)/glycerol solution to produce PTX loaded poly(L-lactic acid) (PLLA) microspheres. 42 It is of the utmost importance to consider its characteristics in droplet and emulsion productions. Hydrophilic PDMS changes to hydrophobic in contact with organic solutions such as dichloromethane. To prevent this phenomenon which circumvents droplet and particle formation, Xu et al. used T-junction ow-focusing geometry. In the conguration, water is in contact with MPs wall which surrenders the dichloromethane phase. Bupivacaine loaded PLGA MPs in the range of 10-50 mm size produced by capillary device release drug more slowly and also have signicantly smaller initial burst release in comparison with conventionally produced particles. 39 Various studies have been conducted to investigate the microuidics capability in the control of MPs shape and morphology. In this regard, PLGA-based MPs fabricated with various shape such as honeycomb, [105][106][107] Golf-ball, 106 snowman microcapsules, 108 porous microbeads. 109 Hussein et al. reported that surface texture is easily controllable using polymers with various hydrophobicity in a microuidics system. 62 They prepared PLGA 100k blend with PLGA 50k -b-PEG 5k /PLGA 100k or PLGA 10k -b-PEG 20k /PLGA 100k and PTX as a solution (10 mg mL À1 ) in DCM as a dispersed phase in the aqueous phase (5 mg mL À1 PVA) with varying ratios in a glass capillary. Results indicated that surface texture is controllable with the blend ratio and consequently it affects encapsulation efficiency and releases kinetics. For example, with an increase in the PEG content (0 to more than 60 wt%) particles surface changes from a smooth to bumpy appearance and nally breaks into nanoscale micelles. Moreover, encapsulation efficiency drops with an increase in the PEG content, e.g. encapsulation of PTX is reported about 92% for neat PLGA 100k (PEG ¼ 0), while in the case of neat PLGA 50k -b-PEG 5k it is about 64%. The authors attributed the results to the fact that higher hydrophobic content of polymer promotes more PTX encapsulation. PEG content also enhances the PTX release, as the highest drug release prole (Fig. 7I) is observed for 80% of PLGA 50k -b-PEG 5k and the lowest for neat PLGA 100k . On one side, the more the hydrophilic content, the more the water adsorption and on the other side the more the PEG content, the higher the roughness and consequently, the more interfacial area available for water diffusion. All together promote drug release for particles with a higher amount of PEG in the polymer blend.
From another point of view, microuidics provides advantages to load agents with various properties in polymeric matrixes. It enables inorganic material loading into PLGA MPs and hybrid MPs production. For instances, the cross-junction microuidic system has been applied to encapsulate CdSe/ ZnS QDs in PLGA MPs with the size of 180 to 550 mm. PLGA along with CdSe/ZnS QDs (4 AE 0.5 nm) in chloroform produces droplets in an aqueous phase containing PVA (1 wt/v%) and chitosan (0.5 wt/v%) in a ow focusing microchip. 110 W/O/W double emulsion template produced in a capillary micro-uidics consisted of an inner aqueous phase containing 10 wt% PEG with 8-hydroxyl-1,3,6-pyrenetrisulfonic acid trisodium salt (a green dye), and sulforhodamine B (a red dye), middle oil phase containing PEG-b-PLA with poly(N-isopropylacrylamide) (PNIPAM)-b-PLGA (as thermosensitive polymer) and dodecylthiol-stabilized gold NPs (a photosensitive agent) in a mixture of chloroform and hexane, outer aqueous solution containing 10 wt% PVA. Aer evaporation, polymersome produced with a double polymer layer (Fig. 7III) in which gold NPs are entrapped in the PNIPAM-b-PLGA part. By examination and optimization with various amounts (2, 5, and 10 wt%) of PNIPAM-b-PLGA, they concluded that 5 wt% is needed to produce thermosensitive polymersome. On the other hand, a higher amount of thermosensitive polymer results in defects and nally ruptures of the polymersome. The release mechanism of polymersome induced by temperature and laser illumination indicated that in thermoresponsive release pores form in the polymersome and releases the load gradually. However, laser triggered release starts with local hot spots formation which nally results in the rupture of bilayer. 111 PLGA microspheres containing TiO 2 NPs on the surface produced in FF chip. PLGA and titanium tetraisopropoxide (TTIP) in DCM generated droplets in aqueous solution 90 wt% of glycerol and 0.5 wt% of PVA. TiO 2 forms upon the contact with aqueous phase as a result of TTIP hydrolysis also make MPs with a wrinkled surface. With the increase in the mass ratio of TTIP/ PLGA (4/30 to 8/30), surface wrinkles get deeper and aer 12/30 particle changes to non-spherical. Tanshinone IIA incorporated into MPs as a model drug with encapsulation efficiency higher than 80% and results from in vitro drug release showed the deeper the wrinkles on the surface the higher drug release rate. 93

Core-shell microparticles
Core-shell MPs or microcapsules contain gas, liquid or solid core (or multiple cores) covered up in the shell. 25,112 These morphologies offer a broad range of advantages such as dual or multiple drug delivery, controlled and prolonged release, protection against active agents. 113 Moreover, the core-shell structure can deliver chemically active agents and protect them from body proteins, immune system, and degradation. Microchannel-based synthesis offers unique opportunities to produce core-shell, multiple cores in one shell and even multiple shells on one core. It is done through solidication of single/multiple emulsions.
For instance, Martín-Banderas et al. encapsulated gentamycin sulfate (GS) in PLGA MPs with one or more core and microcapsules. 114 They used an FF conguration with PLGA and drug solution focused with an aqueous solution to produce MPs. They used the same conguration with multiple concentric needles with air in the inner needle to produce microcapsules. They reported high drug loading (10 to 30%) and encapsulation (42 to 85%) efficiencies for the microcapsules compared with MPs. In another report, alginate shell on PLGA MPs synthesized in a ow focusing capillary microuidic used to control the release of rifampicin in the size range between 15 to 50 mm. The core-shell morphology produced from W/O/W double emulsion templates in which consisted of inner aqueous phase (0.5% sodium alginate and 10% (w/v) PVA), middle oil phase (DCM with 20% PLGA) and outer aqueous phase (10% (w/v) PVA and 4% calcium chloride (CaCl 2 )). Both shell and core sizes affect the drug release and by an increase in the MPs size, initial burst release decreases. Moreover, coreshell structure exhibited higher drug content ($6.4%) than microspheres (4.26 AE 0.54) and also higher encapsulation efficiency (70.47 AE 1.85%) compared with microspheres (46.78 AE 5.89) with similar size ($50 mm). In agreement with results from other investigations, 65 the authors attributed the results to shell layer that prevents the drug from diffusion to solvent in the evaporation stage of the fabrication process. 113,115 PLGA-based MPs containing liquid cores are thermodynamically unstable and shell rupture takes place during solid-ication and degradation that causes to burst release of the loaded drug. Various strategies have been used to overcome this shortcoming and improve MPs stability. For instance, gelatin methacrylate (GelMa) used as a crosslinking agent to avoid rupture or fusion of cores. 43 PLGA-based core-shell structure containing DOX hydrochloride and camptothecin (CPT) fabricated by one-step co-ow capillary microuidics (Fig. 7IV). Double W/O/W emulsion produced in three-concentric capillary tubes. Inner aqueous phase contained hydrophilic drug (DOX hydrochloride) and GelMa (15% w/v) while the outer aqueous phase contains PVA (2% w/v). Hydrophobic drug (CPT) dissolved in DCM with PLGA polymer as a middle oil solution. Emulsions polymerized by UV light to solidify core to produce stable core and release two or more therapeutics sustainably. They achieved the varying size of MPs and core numbers (one, double and triple cores in one shell) with various orice size and uid ow rates. Results illustrated that increase in the shell thickness (22,40, and 60 mm) leads to higher encapsulation efficiency of CPT ($46, 57 and 61%) and DOX hydrochloride (85, 89 and 93%) and higher drug content (4.06 AE 0.02, 6.17 AE 0.15 and 6.88 AE 0.24). In vitro evaluation with human colon cancer cell line (HCT116) (Fig. 7V) and liver cancer cell line (HepG2) revealed the synergistic antitumor effect of two drugs. Antitumor effect of dual drug loaded MPs (less than 20% HCT116 cells and 10% HepG2 cells survive) is higher than individual drugs either DOX or CPT (50% HCT116 cells and 60% HepG2 cells are killed). In another approach, Montazeri et al. 92 improved the PDMS based double ow focusing chip to produce a partially hydrophilic-hydrophobic microuidic device. They added a surfactant (Silwet L-77®) to the curing agent and prepolymer of the PDMS in chip preparation stage to improve the wettability of the chip. They investigated contact angle of water droplet by the surface of PDMS that led them to choose 0.5 wt% among the various concentration of surfactant (0, 0.2, 0.5 and 0.8 wt%) to produce PLGA based MPs with average 20 mm. Results reveal that modication enables H 2 O 2 solution delivery as an oxygen generator into islet transplantation over an extended time up to 30 days.

Janus microparticles
Janus or bifacial MPs have excellent properties such as tunable and controllable asymmetry in shape, composition, the capability to load multiple and even incompatible agents. 89,116 In addition to these properties, recently Janus MPs production in microuidics attracted more attention because of their facility and capability in the synthesis and control of physical and chemical features. 116,117 In order to achieve Janus MPs, two-face droplets are necessary which will solidify to MPs aer droplet consolidation. Therefore, two miscible uids form droplets in a third immiscible uid. Min et al. dissolved PLGA and poly(butyl methacrylate-co-(2-dimethylaminoethyl) methacrylate-co-methyl methacrylate) (p(BMA-co-DAMA-co-MMA)) in an organic solvent (chloroform) as the dispersed phase and the aqueous phase of 10 w/w PVA and 0.1 M NaOH as a continuous phase to produce droplets in a glass capillary. 118 At pH 8.5, by the diffusion of chloroform into the continuous phase and inversion of polymeric domains within the droplets, PLGA core forms in the p(BMA-co-DAMA-co-MMA) shell. Nile red (hydrophobic) and coumarin (hydrophilic) model drugs tend to accumulate in PLGA core and p(BMA-co-DAMA-co-MMA) shell, respectively. However, in pH 10, particles tend to form acorn-shaped Janus MPs (Fig. 7VI). Shell over multiple cores forms when the solvent is replaced with toluene over pH 8.5 to 10. Since the evaporation of toluene happens very fast (2.4-fold faster than chloroform), coalescence of PLGA domain stay separately during evaporation.
PLGA and poly(3-caprolactone) (PCL) MPs with patchy and Janus structures have been produced in droplet microuidics in order to control drug release. 119,120 A solution of PLGA and PCL in DMC or DCM produced organic droplets in a continuous aqueous solution of PVA (2 wt%). The strategy to achieve various morphologies was switching between two organic solvents. Polymer solution in DMC produces MPs with Janus particles (average size 24.42 mm) while in DCM, core-shell (average size 47 mm) structure with PLGA in the core and smooth surface. 119 They investigated the effect of solvent concentration and the mass ratio of PLGA : PCL on the nal MPs morphology and provided a diagram in order to easily choose the ratio for the desired nal product. For example, core-shell MPs generated with a higher portion of PLGA (PLGA : PCL as 7 : 3 or 5 : 5) and anisotropic patchy particles harvested in the lower portions (PLGA : PCL as 3 : 7 or 2 : 8). At an equal mass ratio of polymers, Janus particle and patchy morphology produced with a change in the DCM : DMC volume ratio 1 : 5 and 2 : 1, respectively. 120 In another strategy, Kang et al. 117 produced PLGA Janus particles in which both parts are PLGA. For that purpose, they used two solvents to produce MPs; ethyl acetate (EA) and silicone oil are good and bad solvents for PLGA, respectively. The O/W emulsion in glass capillary which consisted of PLGA dissolved in two solvents as oil phase while continuous PVA aqueous phase as water ow. Faster evaporation of EA compared with removal of silicon oil, in the MPs formation stage produces Janus shape. Researchers claimed that the diameter of each part Janus PLGA MPs is predictable using theoretical and mathematical calculations.

Opportunities and challenges
PLGA DDSs production has started since two decades ago. It's obvious the method and produced NPs have undergone signicant improvements since the rst production of PLGA based NPs in microuidics. PLGA based NPs are being synthesized with advanced features such as targeting ligands, lipid layers, stimuli-responsive, co-loaded with drugs and imaging agents. Various microuidic chips (e.g. 2D, 3D, arc, and origami) have been used to investigate effective factors in the production process such as ow ratio, mixing time and microchannel size. Droplet-based microuidics is being employed for the synthesis of complex MPs. This category of microuidics has offered a powerful platform to control size, size distribution, rapid processing and uploading of drugs with varying hydrophobicity and properties. MPs are being fabricated with core-shell structure and even with multiple cores and/or shell. Apart from the encapsulation of incompatible drugs and/or unstable agents with varying physicochemical properties, the core-shell structure can make it possible to achieve sustained and even sequential release.
Although the technology provides a wide range of opportunities, it is not free from limitations and drawbacks that can hinder its application in large scales.
The rst important issue that should be addressed by prospective researches is the fouling of PLGA NPs during the precipitation process and clogging the micro-size channels. As a matter of fact, the microuidically production of PLGA DDS is considered as a continuous process which avoids batch-tobatch variation. This feature has been arguably accepted as one of the successful aspects of the technology. However, scientists are struggling with a huge number of failed chips in the lab due to the clogging of the microchannels. Not only fouling of NPs but also problems are being reported related to the wettability and hydrophobicity issues in MPs production. Despite all attempts done to solve the problems, results are not admissible and show that microuidics hardware and choice needs to be selected according to every synthesis process.
Moreover, the contribution of various mixers like Tesla micromixers has been introduced and examined before and/or aer focusing section of the synthesis process in continuous microuidics. Two-stage microuidically synthesis has been reported to control the core and shell more precisely in the production inorganic NPs. 121,122 Such multi-stage could be advantageous in the production of PLGA-based DDSs with sophisticated features. Moreover, screening and preclinical test of drugs in living cells are crucial steps in drug discovery, direct purication and cell treatment on the microuidic platform could be considered as additional steps toward the increased efficacy and speed.
Furthermore, manipulation of small volumes of liquids in micro-size has been mentioned as an advantage of these apparatuses; however, it is a double-edged sword and could be a disadvantage at the same time. Low volume of liquids can cause a low throughput of the fabrication process that cannot meet the industrial and large-scale production demand.
Although there are few reports of high production rates suitable for preclinical and clinical demands, researchers need to search a solution for this very crucial issue. Future microuidics should be able to produce a higher amount of drugs without sacricing accuracy and efficiency.

Conclusion
PLGA based drug delivery systems are being produced with various fabrication methods and nowadays there are plenty of them approved by FDA and also many ongoing preclinical and clinical types of research to make their way to industry. For this purpose, researchers from all around the world look for new routes to produce DDS with more sophisticated features to promote the production and also delivery efficiency. In this context, we accented various type of microuidic systems used for the production of PLGA based drug delivery, properties, and applications of PLGA NPs and MPs fabricated by this technology.
All together and in spite of drawbacks, inspiring abilities of the technique hold a great potential to bring and add exciting features to drug delivery systems e.g. controlled release, regulated surface and exibility, stimuli-responsive and etc. knowledge gained from numerous examples of represented PLGA based drug delivery systems prepared in the microuidics can help researchers to select proper reactants, microuidics type and process to fabricate their goal drug delivery system.