Encapsulation of Human IgG Antibody in PLGA Microparticles Fabricated by Coaxial Electrospray

1. Introduction

Cancer immunotherapy is a powerful treatment that activates the immune system of the patient to combat the tumor. Systemic administration of immunotherapy can result in adverse off-target effects, so sustained delivery of immunotherapy following local administration is a promising alternative [1]. Immune checkpoint blockade is a form of immunotherapy that employs monoclonal antibodies such as anti-programmed cell death protein 1 and anti-programmed cell death protein-ligand 1 (anti-PD-1/anti-PD-L1). Current studies have shown promise in controlled-release antibodies for tumor immune checkpoint blockade [2,3]. However, antibody encapsulation technology is new and antibody payload release dynamics have not been fully explored. Recently developed antibody delivery systems and delivery technologies for the immunotherapy of cancer are reviewed here [4,5].

Therapeutic reservoir and drug-polymer matrix systems can extend the delivery of therapies and reduce risk of dose toxicity [6,7]. Typically, these drug-delivery systems consist of a drug payload encapsulated within a polymeric carrier, and the delivery mechanism(s) by which the drug payload is released. Release mechanisms are categorized via payload diffusion, payload-carrier disintegration, polymer swelling, stimuli responsiveness (e.g., pH-activated release), and polymer erosion and degradation [7].

Poly(lactic-co-glycolic acid), or PLGA, is an FDA approved biodegradable copolymer with extensive applications as a drug-delivery carrier [8,9,10]. PLGA is composed of lactic acid and glycolic acid monomers. It is characterized by its molecular weight, monomer ratios, and end groups. PLGA degrades by hydrolysis of the ester bonds of its monomers. It is an attractive polymer for drug-delivery because its molecular weight and monomer ratio can be tuned to modify payload release dynamics. The hydrolysis products are nontoxic and excreted in mammalian systems.

One challenge of many PLGA-based carriers is the tendency to initially release a disproportionately large fraction of the drug cargo [11,12]. This phenomenon is called ‘burst release’ and is generally disfavored when a uniform, sustained release profile is desired. Additionally, challenges in the encapsulation of antibodies in microparticles are due, in part, to the hydrophilic nature of the payload and hydrophobicity of the encapsulating polymer(s). Strategies for improving the encapsulation of small hydrophilic molecules in PLGA microparticles have been outlined for emulsion fabrication, spray-drying, and inkjet printing techniques [13].

PLGA-alginate microspheres for hydrophilic protein delivery were reported by Zhai et al, (2015). A double emulsion and solvent evaporation technique was employed to fabricate PLGA and PLGA-alginate microparticles with hydrophilic protein, bovine serum albumin (BSA). The same technique was used to load PLGA and PLGA-alginate microparticles with rabbit anti-laminin antibody, another protein. The incorporation of alginate in the PLGA particles was reported to provide a more sustained release profile compared to PLGA particles, though PLGA-alginate particles had a reduced biocompatibility [14]. In a separate study, protein-loaded PLGA-alginate particles were fabricated with a water-in-oil emulsification and external gelation. The process parameters, including alginate concentration, cross-linking time, and drying time of BSA-loaded PLGA-alginate particles, were optimized to provide a 13% reduction of the BSA burst-release [15]. However, conventional emulsification techniques risk exposure of the biomolecule payload to organic solvents during fabrication with the potential for loss of bioactivity.

The electrohydrodynamic atomization technique (EHDA), or electrospray, has been used to fabricate single-layer polymeric microstructures from particles and fibers to films at a range of scales (nano to micro) for drug-delivery [12]. A polymeric solution is pumped through a steel needle and a voltage is applied. In the presence of a high voltage, the solution droplet at the tip of the needle deforms into a Taylor cone from which a microjet emerges. As the electric charge builds up on the surface of the polymeric solution and overcomes the Rayleigh limit, the solution breaks off into microdroplets. The solvent evaporates from the microdroplets descend, and microparticles are collected on a ground collector. --

Coaxial electrospray is an extension of the electrospray technique in which a core payload solution is dispensed to an inner capillary and the shell solution is dispensed to an outer capillary. Complex hydrophilic-in-hydrophobic and multi-layer microstructures have been manufactured using this technique [16]. Microparticles have been produced using coaxial electrospray with poly(L-lactic acid) (PLA) with hydrophilic drug core and PLGA with hydrophobic drug shell in dichloromethane (DCM) and ethyl acetate solvents, respectively [16]. The inverse core-shell configuration was also fabricated. A follow-up study showed the encapsulation of two drugs with different degrees of hydrophilicity, paclitaxel and suramin, in PLA and PLGA microspheres to treat mouse U87 glioma [17]. Hydrophilic protein encapsulation in microparticles and fibers with electrospray and electrospinning is reviewed here [18].

The optimization of electrospray parameters can facilitate the fabrication of monodisperse and microstructures [19]. For example, the polymer concentration in the shell solution affects the size and thickness of the particles fabricated. Similarly, the shell flow rate impacts the particle size and shell thickness. The effect of an increased shell flow rate on release profile of curcumin has been shown to reduce burst release and delay release by increasing microparticle shell thickness [20].

Since particle fabrication with electrospray requires the presence of an electric field, a net charge of the drug payload solution can affect its localization within the particle. This is important to consider towards encapsulating proteins and antibodies with this technology. In an electrospinning study, BSA was localized in the core of polyvinyl alcohol (PVA) fibers when dissolved at its isoelectric point [21]. In another study, antibody-encapsulation of bevacizumab in poly-caprolactone (PCL) fibers was achieved using coaxial electrospinning for sustained-release treatment of age-related macular degeneration. The authors showed a first-order antibody release profile with t_1/2 of 11 days from PCL nanofibers fabricated at the commercial pH and transmission electron microscopy (TEM) images suggested antibody localization in the shell. Comparatively, a zero-order release profile with a t_1/2 of 52 days was achieved at the antibody’s isoelectric point and TEM images suggest antibody localization in the core [22].

To date, coaxial electrospray has not been used to fabricate antibody-loaded PLGA microparticles. Thus, there is a gap in the knowledge of electrospray parameters which can achieve antibody encapsulation in PLGA microparticles and characterization of the resultant antibody release profiles. In this manuscript, we focus on the effect of shell flow rate on antibody encapsulation and antibody release-profile which has not been reported previously in the literature. Fluorescent human immunoglobulin-G (IgG-FITC) antibody was used because it is an isotype of various antibodies used in immune checkpoint blockade. We present the development of human antibody-loaded PLGA microparticles using coaxial electrospray at the antibody’s isoelectric point.

2. Materials and Methods

2.1. Materials

Resomer® RG 504 H, Poly(D,L-lactide-co-glycolide) (PLGA) (L:G 50:50, M_w: 38-54 kDa, acid-terminated), immunoglobulin-G (IgG)-FITC from human serum (20 mg/mL), ethyl acetate (ACS reagent, 99.5%), Trizma® Pre-set crystals pH 8.5, phosphate buffered saline (PBS) pH 7.4, Amicon® Ultra Centrifugal Filter (30 kDa MWCO, 2 mL), and poly(vinyl alcohol) (PVA) M_w: 9,000-10,000), were purchased from Millipore Sigma.

2.2. Isoelectric Point of Human Immunoglobulin-G1

The amino acid sequence of the IgG1 antibody was retrieved from the DrugBank database, and the isoelectric point, pI, was calculated using the SIB Swiss Institute of Bioinformatics pI M_w tool. The detailed output can be found in Supplementary Figure S1.

2.2. Preparation of Solutions for Coaxial Electrospray

The concentrations of the solutions are summarized in Table 1. 50 mg/mL of PLGA was dissolved in ethyl acetate for the shell solution. To prepare the core solution (Figure 1a), 1 mL of IgG-FITC from human serum (20 mg/mL) was placed into an Amicon® Ultra Centrifugal Filter (30 kDa MWCO) and centrifuged for 30 minutes to extract the PBS solvent at pH 7.4. Separately, Trizma® pre-set crystals were dissolved in DI water to achieve the pI buffer with pH 8.5. The antibody recovered from centrifugation was diluted to a concentration of 10 mg/mL using the pI buffer.

2.3. Fabrication of Antibody-Loaded PLGA Microparticles with Coaxial Electrospray

The parameters used to fabricate the antibody-loaded microparticles are summarized in Table 1 and the coaxial electrospray system is shown in Figure 1b. Individual syringe pumps were used to dispense the inner and outer solutions at 0.05 mL/hr and 0.5-2 mL/hr, respectively. The solutions were pumped into a 23/17 G coaxial steel needle. A voltage difference of 8.5 kV was applied between the needle and the collector. The microparticles were sprayed onto a ground collector located 15 cm below the needle tip. The microparticles were collected in a stirred deionized (DI) water bath containing 1% PVA. The antibody-loaded microparticles recovered from the DI water-PVA bath were kept atn 4°C for 12 hours for shell hardening, then washed three times by centrifugation in DI water containing 1% PVA. The particles were lyophilized for 24 h using BUCHI Lyovapor™ L-200 Freeze Dryer. Dried particles were stored at –4 °C prior to characterization and release kinetics analysis.

2.4. Characterization of Antibody-Loaded PLGA Microparticles

5 μL of microparticles in DI water were pipetted onto a glass slide and imaged using a Keyence All-in-one microscope. Brightfield and fluorescence images were collected at 40x to investigate colocalization of the microparticles with the fluorescent antibody. Microparticle morphology was imaged with a Zeiss Merlin scanning electron microscope (SEM). Size distribution of the microparticles captured by SEM was determined using ImageJ. Specifically, the circumference (C) and diameters of the (D_M) major and minor axes (D_m) of 50 particles from each preparation were manually identified. Particle zeta potential was analyzed using a Malvern Panalytical Zetasizer Advance Range.

2.5. Characterization of Antibody Release Profile from Antibody-Loaded PLGA Microparticles

In vitro release kinetics were measured by incubating microparticles at 37°C in 1 mL of 1% PBS/BSA buffer solution with a pH of 7.4 (Figure 2). At each time point, the particle suspension was centrifuged at 10x g for 10 minutes at 10°C. Supernatant was extracted and its fluorescence was analyzed relative to an IgG-FITC standard curve using a Tecan Infinite M1000Pro plate reader to calculate the degree of antibody release by the microparticles. After extraction, the supernatant was replaced with 1 mL of 1% PBS/BSA buffer, and the tube was vortexed to resuspend the particles before it was placed back in the incubator.

3. Results

3.1. Morphology and Fluorescence of Antibody-Loaded PLGA Microparticles

Figure 3 is a set of SEM images of microparticles fabricated with the coaxial electrospray method showing non-spherical, round, and sunken particle structure for Formulation 1 (Figure 3a), and discoidal and round particle structure for Formulation 2 (Figure 3b). Optical microscope images and fluoresecence images suggest colocalization of particles and antibody fluoresence signal. In Formulation 1, the fluoresence image suggests antibody core localization whereas in Formulation 2, the fluoresence image suggests antibody shell localization. Table 2 summarizes the averages of particle size and zeta potential for each formulation. Formulation 1 demonstrates lower average circumference as well as major and minor axis size than Fomulation 2 (Table 2). Both particles exhibit a similar negative zeta potential (Table 2).

3.2. Encapsulation Efficiency of Antibody-Loaded PLGA Microparticles

Fluoresence measurements of the particles analyzed according to Figure 2 were used to calculate encapsulation efficiency (EE). The EE was calculated as the percent of total mass of IgG-FITC released from the PLGA particles at the end of the experiments divided by the amount of IgG-FITC used in the electrospray fabrication of the particles.

$$EE\left(\mathrm{\%}\right)=\frac{Total antibody released from PLGA particles}{Total antibody used in electrospray fabrication}\times 100\mathrm{\%}$$

Formulation 1 and Formulation 2 had encapsulation efficiencies of 0.04 $\pm 0.01\mathrm{\%}$ and 1.52 $\pm 0.23\mathrm{\%}$, respectively.

3.3. Fluorescence Release Profile of Antibody-Loaded PLGA Microparticles

Figure 4 shows the percent release of IgG-FITC ng per mg of particles. The release profile of Formulation 1 shows a t_1/2 of 36 hours compared to Formulation 2 with a t_1/2 < 3 hours. Antibody-fluoresence signals measured for 28 days indicated a biphasic release profile with a burst-release followed by a sustained-release for both formulations.

4. Discussion

Dynamic light scattering was attempted to measure the size distribution of the antibody-loaded particles. Due to the collapsed and non-spherical conformation of the particles, it was difficult to produce an accurate measurement of the size distribution using this method. ImageJ was used as an alternative and to measure particle circumference. Measurements of fifty particles demonstrated that Fomulation 1 averaged particles with smaller diameter compared to Fomulation 2, consistent with literature on the impact of increased flow rate on microparticle size [19]. Both Formulations 1 and 2 exhibited a similar negative zeta potential.

The SEM images of the microparticles fabricated show a collapsed non-spherical morphology for Formulation 1 and a collapsed and discoidal morphology for Formulation 2. It is likely that the morphology observed in SEM images is a result of aqueous core solvent evaporation prior to complete PLGA shell hardening. The function of particle morphology is application-dependent. Large discoidal particles can target tumor tissue through vascular adhesion. More specifically, discoidal particles are more likely than spherical particles to localize intratumorally through margination in the tumor vasculature [23]. Further studies are needed to investigate if a discoidal particle morphology like the one produced in Formulation 2 could improve localized immune checkpoint blockade delivery in tumors.

Fluorescence images suggest localization of the antibody with the particles. In Formulation 1, it appears that the fluorescence signal is in the core. In Formulation 2, the fluoresence image suggests there is antibody localized in the shell of the particles.

It was hypothesized that coaxial electrospray of IgG-FITC at its isoelectric point with PLGA shell solution would favor IgG-FITC encapsulation in the core of PLGA microparticles. Though the results suggest this was achieved in Formulation 1, it is unclear why the encapsulation efficiency was low. It is also unclear how increased shell flow rate improved encapsulation efficiency in Formulation 2, yet without the core localization anticipated with the isoelectric point.

During Taylor cone formation, the core and shell solutions come into contact. The miscibility of water in the aqueous core with ethyl acetate in the shell influences how the two solvents interact during this time. In the solution flow rates of Formulation 1, there is a higher water:ethyl acetate volume ratio than in Formulation 2. In Formulation 1, it is possible that ethyl acetate in the shell reached water saturation because ethyl acetate is only partially water miscible. In Formulation 2, a higher flow rate of the shell solution reduce the water:ethyl acetate ratio, thereby increasing the amount of water that is able to be absorbed by the ethyl acetate solution. The absorption of water by the shell solution could explain why is antibody solute was localized in the shell region. This may also explain the increased encapsulation efficiency measured in Formulation 2. The aqueous miscibility of shell solvent may be an underapprecated consideration in the development of polymer shell / biomolecule core sustained delivery systems fabricated by coaxial electrospray.

IgG-FITC localization observed in each Formulation is consistent with the measured release profiles. Antibody encapsulated in the core was released more gradually whereas antibody apparently located near the shell demonstated greater burst-release.

Our findings suggest that in the antibody-encapsulation of IgG-FITC in PLGA microparticles using coaxial electrospray, core-shell solvent interactions may influence encapsulation, antibody localization, and release profiles, even at the antibody’s isoelectric point.