Development of drug-loaded polymer microcapsules for treatment of epilepsy

Yu Chen a, Qi Gu ab, Zhilian Yue *a, Jeremy M. Crook acd, Simon E. Moulton e, Mark J. Cook afg and Gordon G. Wallace *a
aARC Centre of Excellence for Electromaterials Science, Intelligent Polymer Research Institute, AIIM Facility, Innovation Campus, University of Wollongong, Northfields Avenue, Wollongong, NSW 2522, Australia. E-mail: zyue@uow.edu.au; gwallace@uow.edu.au
bState Key Laboratory of Stem cell and Reproductive Biology, Institute of Zoology, Chinese Academy of Sciences, 1 Beichen West Road, Chaoyang District, Beijing 100101, P. R. China
cDepartment of Surgery, University of Melbourne, St Vincent's Hospital, 35 Victoria Parade, Fitzroy, Victoria 3065, Australia
dIllawarra Health and Medical Research Institute, University of Wollongong, Wollongong, New South Wales 2522, Australia
eARC Centre of Excellence for Electromaterials Science, Faculty of Science, Engineering and Technology, Swinburne University of Technology, Hawthorn, Victoria 3122, Australia
fClinical Neurosciences, St Vincent's Hospital, 5th Floor, Daly Wing, 35 Victoria Parade, Fitzroy, Victoria 3065, Australia
gDepartment of Medicine, University of Melbourne, St Vincent's Hospital, 35 Victoria Parade, Fitzroy, Victoria 3065, Australia

Received 15th July 2017 , Accepted 7th September 2017

First published on 7th September 2017


Despite significant progress in developing new drugs for seizure control, epilepsy still affects 1% of the global population and is drug-resistant in more than 30% of cases. To improve the therapeutic efficacy of epilepsy medication, a promising approach is to deliver anti-epilepsy drugs directly to affected brain areas using local drug delivery systems. The drug delivery systems must meet a number of criteria, including high drug loading efficiency, biodegradability, neuro-cytocompatibility and predictable drug release profiles. Here we report the development of fibre- and sphere-based microcapsules that exhibit controllable uniform morphologies and drug release profiles as predicted by mathematical modelling. Importantly, both forms of fabricated microcapsules are compatible with human brain derived neural stem cells and differentiated neurons and neuroglia, indicating clinical compliance for neural implantation and therapeutic drug delivery.


Introduction

Approximately 90% of the new drugs that are successful in preclinical studies fail in clinical trials.1 The high failure rate in new drug development is not due to a lack of drug potency, but rather a lack of efficacy because of side effects arising from toxicity, poor pharmacokinetics and pharmacodynamics. A promising strategy for overcoming these drawbacks is to develop an appropriate drug delivery system for targeted and efficient application. For treatment of brain diseases, drug delivery to the central nervous system (CNS) represents a unique challenge. A number of factors, including physical permeability barriers such as the blood–brain barrier (BBB) and the blood-cerebrospinal fluid (blood-CSF) barrier, and expression of multidrug drug efflux transporters at the barriers, work collectively to restrict the entry of many pharmaceuticals into the brain, causing drug resistance.2–6 In the case of epilepsy, 30–40% of patients remain drug resistant with poor clinical outcomes, and this represents a significant hurdle for therapy of intractable epilepsy.7

In order to deliver drugs to the brain, considerable effort has been made to develop nanocarriers to aid in systemic delivery.8–15 However, local drug delivery directly to the site of action has the potential to significantly improve the therapeutic efficacy of epilepsy medication. It directly bypasses the brains physical barriers, which may result in more efficient drug delivery and bio-distribution. Consequently a lower therapeutic dosage may be required, with concomitant reduction in side effects.16–18 These advantages have prompted research and development of local drug delivery systems for treatment of epilepsy.19–22 For example, Williamson et al. developed an organic electronic ion pump, composed of poly(3,4-ethylenedioxythiophene) doped with polystyrene sulfonate electrodes, for on-demand and site-specific delivery of an inhibitory neurotransmitter, gamma-aminobutyric acid (GABA). The delivery of GABA was shown to result in quick and localized suppression of epileptiform activity.22 Pritchard et al. developed silk fibroin coatings on solid reservoirs of the anticonvulsant adenosine. By modulating the coating thickness and degree of crystallinity, sustained release of adenosine with various release profiles including zero order release profiles was achieved over a period of two weeks.20

We have developed levetiracetam-loaded PLGA sheets. Following implantation in the tetanus toxin model of temporal lobe epilepsy in rats, the polymer sheets have shown to significantly shorter seizures and a trend towards fewer seizures for up to 1 week.23 More recently, we have developed lacosamide-loaded electrospun poly(D,L-lactic-co-glycolic acid) (PLGA) nanofibers, which could affect seizure activity in a rat model of absence epilepsy for up to 7 weeks.17 In addition, using a conventional emulsion method, we have developed injectable phenytoin loaded polymeric microspheres.24 A single injection of the microspheres into the hippocampus of a rat tetanus toxin model of temporal lobe epilepsy can control seizures for the expected period of time that is in accord with in vitro release data. These studies have demonstrated the potential of local drug delivery systems for treatment of epilepsy. In the meantime, they have also highlighted the need for further development of local drug delivery systems that exhibit controllable uniform morphologies, high drug loading efficiency and sustained and predictable drug release profiles. This has prompted us to conduct the current study.

In this paper, we present a simple electrojetting (electrospinning and electrospraying) technique to fabricate drug-laden microcapsules for sustained delivery of antiepilepsy drugs (AEDs). Lacosamide was selected as a model drug, and biodegradable poly(D,L-lactic-co-glycolic acid) (PLGA) employed as the carrier material due to its biocompatibility with brain tissue.25–27 The as-fabricated microcapsules include microspheres, flattened (oblate) microspheres (spheroids), and microfibers, which can be utilized as injectable/implantable systems for site-specific delivery of AEDs. They all have uniform morphology, controllable size distribution, enable high efficiency drug encapsulation, and release profiles that are consistent with mathematical modelling for predictability. Importantly, the microcapsules are compatible with human frontal cortical neural stem cells (NSCs) and derivative neurons and neuroglia, indicative of clinical compliance for therapeutic use.

Materials and methods

Materials

The anti-epilepsy drug, lacosamide, was provided by UCB Pharma Pty Ltd. Poly(D,L-lactic-co-glycolic acid) (Mw ∼ 60[thin space (1/6-em)]000 Da, with a 75/25 molar ratio of lactide to glycolide) (PLGA 75/25) was purchased from Purac, Singapore. Chloroform, methanol, and acetonitrile were all analytical grade from Sigma-Aldrich. All the others chemicals and reagents were purchased from Sigma-Aldrich and used as received.

Microcapsule fabrication via electrojetting

The microcapsules, including microspheres and microfibers, were fabricated using NANON-01A electrospinning system (MECC Co. Ltd, Japan) at ambient temperature. A range of PLGA/lacosamide (w/w, 10/1) solutions were prepared in chloroform, with the PLGA concentration varying from 1.5 wt%, 4.5 wt%, to 14.0 wt%, respectively. For electrojetting, each solution was loaded into a plastic syringe equipped with a 23 gauge stainless steel needle. The distance between the tip of the needle and the aluminium foil was 12 cm, and the voltage was 10 kV for electrospraying and 21 kV for electrospinning. The feed rate was controlled at 0.5 mL h−1 using a syringe pump, and the electrojetted samples were collected using aluminium foil. The samples were then further dried in a vacuum oven at room temperature for 48 hours to remove any residual organic solvent.

Morphological and dimensional statistical analysis

A field emission scanning electron microscope (FESEM, JEOL JSM-7500FA) was used to examine the morphologies of the as-fabricated microcapsules. The samples were sputter-coated with 20 nm gold prior to SEM testing. Dimensional statistical analysis was conducted by quantitatively evaluating the high-magnification SEM micrographs using the imaging software, Leica Application Suite. For microcapsules including microspheroids and microspheres, aspect ratio is defined as the ratio of the length (l) to the width (w) of a microcapsule, and diameter is defined as the length of a microcapsule.

In vitro drug release study

Drug release study was conducted under sink condition in artificial cerebrospinal fluid (aCSF) at 37 °C in a shaking water bath. An appropriate amount of solid microspheroids (2.47 ± 0.88 m), microspheres (3.81 ± 0.38 mg) and microfibers (4.43 ± 0.11 mg) were used in this release study. Each sample was immersed in 1 mL of aCSF, and the released solution was removed at various time points and replaced with 1 mL of fresh aCSF. All samples were kept at −20 °C, before being analysed by high performance liquid chromatography (HPLC). The details of the HPLC testing method were reported in our previous work.21 In addition, to evaluate the drug encapsulation efficiency, an extraction method was used as previously reported.28 Briefly, the fabricated microspheroids, microspheres or microfibers were extracted with methanol. The methanol extract was then quantified by HPLC to determine the actual drug loading in each group of fabricated microcapsules.

Mathematical modelling of the release profiles of the microcapsules

The release profiles of the microcapsules were simulated using Fick's second law of diffusion subject to appropriate boundary conditions.29,30 When the surface resistance to mass transfer at the surface is negligible, the fraction of the drug released from the microcapsules (Mt/M) at any time (t) can be expressed as follows:
 
image file: c7bm00623c-t1.tif(1)
where, d is the diameter of the microcapsules, De is the effective diffusivity of drug, i = (2, 3, 3) and λn = (nπ) for microspheroids and microspheres, λn = (2.41, 5.52, 8.65, 11.8, 14.93…) for microfiber.31 The mathematical modelling study is based on eqn (1), using MATLAB Curve Fitting Toolbox.

NSC culture and differentiation

Working stocks of human NSCs (ReNcell CX, SCC007, Millipore) were maintained under 5% CO2 at 37 °C, seeding at a density of 2–3 × 106 cells in proliferation medium comprising NeuroCult NS-A (#5751, Stem Cell Technologies) with 2 μg mL−1 heparin, 20 ng mL−1 FGF2 and 20 ng mL−1 EGF (AF-100-15, Peprotech) on laminin (L6274, Life Technologies) coated 6-well plates (Greiner Bio-One). Cells were passaged at confluency every 5–7 days by digesting in TrypLE (Life Technologies) for 3 min at 37 °C. Differentiation of NSCs was performed 3 days after initially seeding in proliferation medium, using neural differentiation medium comprising two parts DMEM/F-12 (11330-032, Life Technologies), one part Neurobasal (21103-049, Life Technologies) supplemented with 0.5% N2 (17502048, Gibco) and 50 ng mL−1 brain derived neurotrophic factor (BDNF, 450-02, Peprotech) for up to 10 days.

NSC viability analysis

PrestoBlue™ cell viability reagent was used for NSC viability studies, according to the manufacturer's instructions. Briefly, cells were incubated with the reagent in culture medium for 1 h at 37 °C. Following incubation, for each sample, 100 μL supernatant was transferred to a well of a 96-well plate and screened by a microplate reader (POLARstar Omega) to read fluorescence intensity. After processing, samples were rinsed in culture medium and returned to culture, with the process repeated for each time point until the study was completed.

NSC immunocytochemistry

Samples were fixed with 3.7% PFA solution in PBS at RT for 30 min, rinsed in PBS, and then blocked and permeabilized overnight at 37 °C with 5% (v/v) donkey serum in PBS containing 0.3% (v/v) Triton X-100 (Sigma). Samples were subsequently incubated with fluorescence conjugated antibodies glial fibrillary acid protein (GFAP; mouse, 1[thin space (1/6-em)]:[thin space (1/6-em)]1000; Millipore), mature neuronal microtubule–associated protein 2 (MAP2; mouse, 1[thin space (1/6-em)]:[thin space (1/6-em)]100; Sigma) and early neuronal-specific class III beta-tubulin protein (TUJ1; mouse, 1[thin space (1/6-em)]:[thin space (1/6-em)]100; Invitrogen), to label neuroglia (including astrocytes), and mature and newly generated neurons, respectively. On the second day, samples were rinsed with 0.1% Triton X-100 in PBS three times, and samples with unconjugated primary antibody were incubated with Alexa Fluor tagged secondary antibody (1[thin space (1/6-em)]:[thin space (1/6-em)]1000; Invitrogen) for 1 h at 37 °C. Nuclei were visualised with 4′,6-diamidino-2-phenylindole (DAPI, 10 μg ml−1) at RT for 10 min and antifade reagent (Invitrogen) was employed to preserve fluorescence signal. Samples were mounted onto glass coverslips using Aquamount (ThermoScientific) and imaged on a confocal microscope (Leica TSC SP5 II). Images were collected and analysed using Leica Application Suite AF (LAS AF) software (Leica).

Scanning electron microscopy (SEM)

SEM was performed as previously described.32 Briefly, samples were fixed in 3.7% paraformaldehyde (PFA) for 10 min and then serially dehydrated in 30%, 50%, 70%, 85%, 95%, and 100% ethanol before Critical Point Drying using a LeicaEMCPD030 instrument. The samples were then coated with 20 nm platinum followed by imaging with a JEOL JSM-7500FA.

Results and discussion

Fabrication of lacosamide-loaded microcapsules by electrojetting

Electrojetting is an electrohydrodynamic process and is a simple, versatile and cost effective technology that employs an electrically charged jet of a polymer solution to fabricate nano or microscale fibers (i.e., electrospinning) or particles (i.e., electrsopraying).33–38 Electrojetting is governed by the interactions between the electrostatic repulsion induced by an applied electric field, and surface tension of a liquid droplet. At the tip of the capillary, due to the electrostatic repulsion and the surface tension, the hemispherical surface of the polymer droplet is distorted into a conical shape that is known as the Taylor cone. When the electrostatic repulsion surpasses the surface tension, liquid ejection will occur at the surface of the Taylor cone. Electrospinning (Fig. 1a) typically occurs when the polymer concentration and molecular weight are sufficiently high, so that the Taylor cone is stable, and the fluid does not break up into droplets but forms a stable liquid jet. The liquid jet will undergo a whipping or bending motion process, giving rise to the formation of fibres.33 However, when the polymer concentration is below a threshold and/or the polymer molecular weight is low enough, the Taylor cone becomes unstable, and the liquid jet breaks up due to varicose instabilities and hence fine droplets are formed. The electrostatic forces among the droplets enable self-dispersing of the droplets in space with minimal droplet agglomeration. Further evaporation of the solvent leads to concentration and solidification of the droplets, i.e., the formation of polymeric microsphere. This process is also known as electrospraying, which is illustrated in Fig. 1b.39
image file: c7bm00623c-f1.tif
Fig. 1 Schematic illustration of the electrojetting setup, with scanning electron microscope (SEM) images. (a) Electrospinning PLGA/lacosamide microfibers (b) electrospraying PLGA/lacosamide microspheres.

Therefore, under a given electric filed, the difference between the electrospinning and electrospraying lies in the chain entanglement density of the polymer solution.34 In this study, the only variable affecting the chain entanglement density is the PLGA concentration. The final electrojetted structure is mainly determined by the PLGA concentration. By carefully varying the PLGA concentration in the solutions, we have fabricated three kinds of microcapsules, including microspheroids, microspheres, and microfibers.

Fig. 2 shows the SEM micrographs of lacosamide-loaded microspheroids (obtained from 1.5 wt% PLGA and at 10 kV), microspheres (obtained from 4.5 wt% PLGA and at 10 kV), and microfibers (obtained from 14 wt% PLGA and at 21 kV). Electrospraying resulted in the formation of either microspheres or microspheroids, while electrospinning led to the formation of microfibers. These results are consistent with the previous studies, where the polymer concentration is demonstrated to be the most critical parameter in determining the morphology of electrojetted microcapsules.30,34 In the current study, at a low PLGA concentration such as 1.5 wt% and 4.5 wt%, even with an increased applied voltage up to 21 kV, only micro-spherical structures were fabricated.


image file: c7bm00623c-f2.tif
Fig. 2 Various PLGA/lacosamide microcapsules. Scanning electron microscope (SEM) images of (a, b) microspheroids, (c, d) microspheres, (e, f) microfibers.

Dimensional and shape uniformity analysis

Evaluation of the dimensional and shape uniformity was conducted by analysis of the SEM images of the as-prepared microcapsules. As shown in Fig. 3a, the diameters are 3.01 ± 0.51 μm for the microspheroids, 5.10 ± 0.78 μm for the microspheres, and 1.38 ± 0.15 μm for the microfibers, respectively. The diameter distribution of microspheres, microspheroids, and microfibers is shown in Fig. 3b. All the as-fabricated microcapsules demonstrate uniform morphology with narrow size distribution.
image file: c7bm00623c-f3.tif
Fig. 3 (a) Size and (b) size distribution of microspheroids, microspheres, and microfibers. Aspect ratio of PLGA/lacosamide, (c, e, f) microspheroids, and (d, g, h) microspheres; aspect ratio defined as the ratio of the l to w as shown in figures.

For electrospinning, as the PLGA concentration is high as 14 wt%, sufficient polymer entanglements are present and are responsible for the formation of a stable liquid jet of PLGA that undergo continuous stretching during the whipping or bending motion process. Compared to the microcapsules produced by electrospray, the diameters of the electrospun microfibers are much thinner.40

For electrospraying, the dimension and shape of the microcapsules are controlled by two processes; (i) rapid solvent evaporation of the droplets, and (ii) polymer diffusion during evaporation. However, rapid solvent evaporation and polymer diffusion do not necessarily lead to microspheres. When a lower PLGA concentration, such as 1.5 wt%, was used in present work, spheroid morphology is formed. This is probably due to incomplete solvent evaporation, as such the electrosprayed microcapsules are still in liquid form when reaching the collector.41 Increasing the polymer concentration to 4.5 wt% prompts the formation of standard microspheres with increased particle sizes. Since increasing the polymer concentration leads to an increase in the solution viscosity, and a reduction in the surface tension. In accordance with Hartman's study, decreasing the surface tension result in increasing the size of electrospray PLGA microspheres.42

Aspect ratio (AR), expressed as l to w (shown in Fig. 3f and h), is an important parameter to evaluate the shape uniformity of the particles. We measured dimensions of microspheroids and microspheres in two perpendicular directions. The AR of the microspheres is 1.03, and 1.05 for the microspheroids. Both types of microcapsules possess excellent shape uniformity.

In vitro drug release and mathematical modelling study

Drug loading determined by the extraction method is 8.1 ± 0.63 wt% for the microspheroids, 8.36 ± 0.31 wt% for the microspheres, and 8.39 ± 0.19 wt% for the microfibers. As shown in Fig. 4a, all the microcapsules exhibit very high drug loading efficiency; namely 89.1 ± 6.9% for microspheroids, 92.0 ± 3.5% for microspheres, and 92.3 ± 2.1% for microfibers. The encapsulation efficiencies are significant greater than those prepared using other techniques, including emulsion, suspension, and emulsion polymerization,43 solvent evaporation,43,44 spray drying,45 and layer-by-layer encapsulation.46
image file: c7bm00623c-f4.tif
Fig. 4 (a) Drug loading efficiency of PLGA/lacosamide microspheroids, microspheres, and microfibers, (b) cumulative release profiles of various PLGA microcapsules (mean ± standard deviation), (c) comparison of the mathematical modelling results (solid lines) with experimental drug release profiles (symbols).

Fig. 4b shows the in vitro drug release profiles of the microspheroids, microspheres, and microfibers. The release profiles vary significantly with the shape and morphology of the microcapsules. While the microfibers demonstrate the least initial burst release, the spheroids exhibit the most rapid release characteristics, with more than 95% of the lacosamide being eluted within ∼52 hours. Within the same period, the cumulative release of the lacosamide from the microspheres and microfibers is approximately 74% and 55% of the respective total drug loading.

In clinical application, for short term seizure prophylaxis, such as after brain surgery, it's preferable to have rapid release kinetics to reach a therapeutic level. The microfibers are more clinically relevant with longer duration of drug release.47–51 Compared to microspheroids and microspheres, the microfibers have demonstrated the most favourable drug release characteristics. Therefore the fibers would be likely to have the greatest longevity after implantation and be likely to provide the most consistent seizure control as the release rates are more constant compared to the other formulations. An ideal implantable seizure control device should improve the drug release kinetics for a consistent and controllable dosage for the life of the implant.19 The fabrication methods demonstrated here can be readily modified to alter the AED release profile. For example, by applying core–shell spinnerets we have developed core–shell microfibers,17 where the shell structures act as modifiable diffusion barrier to the drug-laden cores.

Typically, drug release from biodegradable microcapsules is controlled by diffusion at an initial stage and then a combination of diffusion and erosion at the later stage.52 During incubation in the release medium, drug diffusion occurs initially within a thin surface layer, and subsequently with water penetration into the bulk of the microcapsules, enhanced diffusive transport of the remaining lacosamide in the microcapsules occurs. Therefore, the morphology and dimension of the PLGA microcapsules have a significant influence on the release of lacosamide. Most of the reported mathematical modelling studies for biodegradable microcapsules drug delivery systems are based on a single, zero-order process; or a process that is governed only by diffusional mass transfer or chemical reactions.53 In order to more accurately predict drug release kinetics for biodegradable microcapsules, mathematical models that take account of the effects of composition and geometry (size and shape) on the drug release are required.

In this study, it is assumed that drug release from the microcapsules is predominantly diffusion-controlled within the period of study. This assumption is based on our and others’ in vitro studies that indicate minimal polymer degradation of the microcapsules within a period of up to three weeks.21,30,53 According to Fick's second law of diffusion, the fraction of the released drug from the microcapsules (Mt/M) at any time (t) can be expressed as eqn (1). In this model, a well-mixed external aqueous phase with a negligibly small drug concentration is assumed. And the drug is assumed to be homogeneously distributed throughout the polymer matrix with an initial drug concentration higher than the solubility of the drug in the polymer. For the microfiber geometry, the drug diffusional release is considered from a cylindrical matrix. The eigenvalues λn, the roots of eigenfunctions are λn = 2.41, 5.52, 8.65, 14.93, and so on.31 The mathematical model expression as:

 
image file: c7bm00623c-t2.tif(2)

For the microspheroid and microsphere geometry, the diffusional drug release is considered from a spherical matrix. The eigenvalues λn, the roots of eigenfunctions are nπ, therefore eqn (1) can be expressed as:

 
image file: c7bm00623c-t3.tif(3)

Fig. 4c shows the least-squares fitting results by applying these mathematical models. The initial drug loading is determined by loading efficiency testing results of HPLC. Our results demonstrated that drug release profiles using the above models can be predicted and agree with the experimental drug release data. The computed effective diffusivities (De) are 3.29 × 10−12 cm2 min−1 for microspheroids, 1.89 × 10−12 cm2 min−1 for microspheres, and 1.23 × 10−13 cm2 min−1 for microfibers, well within the range of values previously reported for hydrophobic drugs in PLGA microcapsules.29 Compared to the spheroids and spheres, the microfibers possess the lowest effective diffusivities (De) during this time-dependent diffusion process.

Human NSC viability and differentiation

Human NSCs are native “adult stem cells” that are capable of self-renewal and differentiation into multi neural lineages including neurons and supporting neuroglia. In this study, we have demonstrated the neuro-cytocompatibility of our electrojetted spheres and fibres using clinically relevant human frontal cortical NSCs and their differentiation to neurons and supporting neuroglia. Neurocompatibility was supported by cell attachment (Fig. 5a) and viability (Fig. 5b) during extended cell culture with fibres and spheres. Growth profiling indicated cell proliferation on all substrates, similar to control (Fig. 5b). Specifically, quantitative measurement of cell viability marker PrestoBlue® supported normal growth kinetics in all cases, with cell proliferation increasing exponentially up to 8 days culture. As expected, cell viability decreased following peak cell growth, reflecting the limits of culture due to media exhaustion.54 These results demonstrated the normal survival of human NSCs on the polymeric substrates. Importantly, differentiation of NSCs resulted in neuronal cells cultured for up to 10 days with densely packed neurites adhered to the fibres (Fig. 5c) and spheres (Fig. 5c). Immunophenotyping confirmed successful differentiation and co-culture of fibres and spheres with pan-neuronal markers, MAP2 and TUJ1 expressing neurons and radial glial and astrocyte marker GFAP expressing glial cells (Fig. 6). The NSC attachments to and ongoing culture on the fibres and spheres indicated cell supportability by the electrospun polymeric structure.55,56 Notwithstanding the value of our findings, it is important to note that our in vitro cell-based modelling involves testing in a controlled environment that does not precisely replicate the cellular conditions within an organism. Therefore, future studies will entail animal-based modelling for in vivo implantation.
image file: c7bm00623c-f5.tif
Fig. 5 Human NSC viability and differentiation on spheres or fibres. NSCs were cultured with (a) fibres and (b) spheres following successful attachment. (c) NSC viability and proliferation kinetics during culture with fibres and spheres (mean ± std.). Morphology of differentiated NSCs during culture with (d, e) fibres and (f, g) spheres compared to (h, i) conventional plate-based culture.

image file: c7bm00623c-f6.tif
Fig. 6 Immunophenotyping of differentiated NSCs. (a, b) DAPI (blue) colocalised with neuronal markers MAP2 (green), and TUJ1 (red), and glial marker GFAP (purple) expressed by differentiated NSCs cultured with fibres. (c, d) DAPI (blue) colocalised with neuronal markers MAP2 (green) and TUJ1 (red), and glial marker GFAP (purple) expressed by differentiated NSCs cultured with spheres. All scale bars: 50 μm.

Conclusion

In conclusion, we have fabricated microstructured drug delivery systems using a simple electrojetting (electrospraying and electrospinning) technology for local delivery of AEDs. These microcapsules have demonstrated a number of advantages, including controllable and uniform morphology and narrow size distribution, high efficiency drug encapsulation, sustained and predictable drug release characteristics, and neuro-cytocompatibility as evidenced by human NSC support and differentiation. Our findings broadly support the utility of the local drug delivery systems for treating human brain disorders such as epilepsy. In addition, the method presented here could be applicable to the fabrication of other microstructured delivery systems using different drugs and/or polymers.

Conflicts of interest

There are no conflicts to declare.

Acknowledgements

Funding from the Australian Research Council Centre of Excellence Scheme (Project Number CE 140100012) is gratefully acknowledged. GGW is grateful to the ARC for support under the Australian Laureate Fellowship scheme (FL110100196). The authors are grateful to UCB Pharma for continued support. The authors gratefully acknowledge Mr Tony Romeo and Mr Mitchell Nancarrow from Electron Microscope Centre (EMC) for technical assistance of SEM images statistical analysis. The authors also gratefully acknowledge the use of facilities within the Australian National Fabrication Facility (ANFF).

References

  1. M. Hay, D. W. Thomas, J. L. Craighead, C. Economides and J. Rosenthal, Nat. Biotechnol., 2014, 32, 40–51 CrossRef CAS PubMed.
  2. N. J. Abbott, L. Ronnback and E. Hansson, Nat. Rev. Neurosci., 2006, 7, 41–53 CrossRef CAS PubMed.
  3. R. Cecchelli, V. Berezowski, S. Lundquist, M. Culot, M. Renftel, M.-P. Dehouck and L. Fenart, Nat. Rev. Drug Discovery, 2007, 6, 650–661 CrossRef CAS PubMed.
  4. N. J. Abbott, E. U. Khan, C. M. Rollinson, A. Reichel, D. Janigro, S. M. Dombrowski, M. S. Dobbie and D. J. Begley, in Drug resistance in epilepsy: the role of the blood-brain barrier, Novartis Foundation Symposium, John Wiley, Chichester, UK, 2002, pp. 38–53 Search PubMed.
  5. B. Engelhardt and L. Sorokin, Semin. Immunopathol., 2009, 31, 497–511 CrossRef PubMed.
  6. C. Martin, A. De Baerdemaeker, J. Poelaert, A. Madder, R. Hoogenboom and S. Ballet, Mater. Today, 2016, 19, 491–502 CrossRef CAS.
  7. M. J. Cook, T. J. O'Brien, S. F. Berkovic, M. Murphy, A. Morokoff, G. Fabinyi, W. D'Souza, R. Yerra, J. Archer, L. Litewka, S. Hosking, P. Lightfoot, V. Ruedebusch, W. D. Sheffield, D. Snyder, K. Leyde and D. Himes, Lancet Neurol., 2013, 12, 563–571 CrossRef PubMed.
  8. S. B. Tiwari and M. M. Amiji, Curr. Drug Delivery, 2006, 3, 219–232 CrossRef CAS.
  9. Y. Miura, T. Takenaka, K. Toh, S. Wu, H. Nishihara, M. R. Kano, Y. Ino, T. Nomoto, Y. Matsumoto, H. Koyama, H. Cabral, N. Nishiyama and K. Kataoka, ACS Nano, 2013, 7, 8583–8592 CrossRef CAS PubMed.
  10. S. Krol, J. Controlled Release, 2012, 164, 145–155 CrossRef CAS PubMed.
  11. Y. Chen and L. Liu, Adv. Drug Delivery Rev., 2012, 64, 640–665 CrossRef CAS PubMed.
  12. J. Mo, L. He, B. Ma and T. Chen, ACS Appl. Mater. Interfaces, 2016, 8, 6811–6825 CAS.
  13. M. Gregori, D. Bertani, E. Cazzaniga, A. Orlando, M. Mauri, A. Bianchi, F. Re, S. Sesana, S. Minniti, M. Francolini, A. Cagnotto, M. Salmona, L. Nardo, D. Salerno, F. Mantegazza, M. Masserini and R. Simonutti, Macromol. Biosci., 2015, 15, 1687–1697 CrossRef CAS PubMed.
  14. L. Wei, X.-Y. Guo, T. Yang, M.-Z. Yu, D.-W. Chen and J.-C. Wang, Int. J. Pharm., 2016, 510, 394–405 CrossRef CAS PubMed.
  15. F. Meng, S. Asghar, Y. Xu, J. Wang, X. Jin, Z. Wang, J. Wang, Q. Ping, J. Zhou and Y. Xiao, Int. J. Pharm., 2016, 506, 46–56 CrossRef CAS PubMed.
  16. S. Mura, J. Nicolas and P. Couvreur, Nat. Mater., 2013, 12, 991–1003 CrossRef CAS PubMed.
  17. S. H. Bauquier, J. L. Jiang, Z. Yue, A. Lai, Y. Chen, S. E. Moulton, K. J. McLean, S. Vogrin, A. J. Halliday and G. Wallace, Int. J. Polym. Sci., 2016 Search PubMed , 6594960.
  18. N. Wang, P. Sun, M. Lv, G. Tong, X. Jin and X. Zhu, Biomater. Sci., 2017, 5, 1041–1050 RSC.
  19. A. J. Halliday, S. E. Moulton, G. G. Wallace and M. J. Cook, Adv. Drug Delivery Rev., 2012, 64, 953–964 CrossRef CAS PubMed.
  20. E. M. Pritchard, C. Szybala, D. Boison and D. L. Kaplan, J. Controlled Release, 2010, 144, 159–167 CrossRef CAS PubMed.
  21. Y. Chen, Z. Yue, S. E. Moulton, P. Hayes, M. J. Cook and G. G. Wallace, J. Mater. Chem. B, 2015, 3, 7255–7261 RSC.
  22. A. Williamson, J. Rivnay, L. Kergoat, A. Jonsson, S. Inal, I. Uguz, M. Ferro, A. Ivanov, T. A. Sjöström, D. T. Simon, M. Berggren, G. G. Malliaras and C. Bernard, Adv. Mater., 2015, 27, 3138–3144 CrossRef CAS PubMed.
  23. A. J. Halliday, T. E. Campbell, T. S. Nelson, K. J. McLean, G. G. Wallace and M. J. Cook, J. Clin. Neurosci., 2013, 20, 148–152 CrossRef CAS PubMed.
  24. J. L. Jiang, Z. Yue, S. H. Bauquier, A. Lai, Y. Chen, K. J. McLean, A. J. Halliday, Y. Sui, S. Moulton and G. G. Wallace, Restor. Neurol. Neurosci., 2015, 33, 823–834 CAS.
  25. P. Menei, V. Daniel, C. Montero-Menei, M. Brouillard, A. Pouplard-Barthelaix and J. P. Benoit, Biomaterials, 1993, 14, 470–478 CrossRef CAS PubMed.
  26. E. Bible, D. Y. S. Chau, M. R. Alexander, J. Price, K. M. Shakesheff and M. Modo, Biomaterials, 2009, 30, 2985–2994 CrossRef CAS PubMed.
  27. H. C. S. Abeysinghe, L. Bokhari, A. Quigley, M. Choolani, J. Chan, G. J. Dusting, J. M. Crook, N. R. Kobayashi and C. L. Roulston, Stem Cell Res. Ther., 2015, 6, 186 CrossRef PubMed.
  28. J. Salonen, L. Laitinen, A. M. Kaukonen, J. Tuura, M. Björkqvist, T. Heikkilä, K. Vähä-Heikkilä, J. Hirvonen and V. P. Lehto, J. Controlled Release, 2005, 108, 362–374 CrossRef CAS PubMed.
  29. D. Y. Arifin, L. Y. Lee and C.-H. Wang, Adv. Drug Delivery Rev., 2006, 58, 1274–1325 CrossRef CAS PubMed.
  30. P. Fattahi, A. Borhan and M. R. Abidian, Adv. Mater., 2013, 25, 4555–4560 CrossRef CAS PubMed.
  31. G. N. Watson, A treatise on the theory of Bessel functions, Cambridge University Press, 1995 Search PubMed.
  32. J. Xie, M. R. MacEwan, A. G. Schwartz and Y. Xia, Nanoscale, 2010, 2, 35–44 RSC.
  33. Y. Dzenis, Science, 2004, 304, 1917–1919 CrossRef CAS PubMed.
  34. S. L. Shenoy, W. D. Bates, H. L. Frisch and G. E. Wnek, Polymer, 2005, 46, 3372–3384 CrossRef CAS.
  35. Y. Chen, D. Han, W. Ouyang, S. Chen, H. Hou, Y. Zhao and H. Fong, Composites, Part B, 2012, 43, 2382–2388 CrossRef CAS.
  36. P. Pal, P. Dadhich, P. K. Srivas, B. Das, D. Maulik and S. Dhara, Biomater. Sci., 2017, 5, 1786–1799 RSC.
  37. G. Xu, Y. Tan, T. Xu, D. Yin, M. Wang, M. Shen, X. Chen, X. Shi and X. Zhu, Biomater. Sci., 2017, 5, 752–761 RSC.
  38. Z. Chen, Z. Chen, A. Zhang, J. Hu, X. Wang and Z. Yang, Biomater. Sci., 2016, 4, 922–932 RSC.
  39. N. Bock, T. R. Dargaville and M. A. Woodruff, Prog. Polym. Sci., 2012, 37, 1510–1551 CrossRef CAS.
  40. J. Zheng, H. Zhang, Z. Zhao and C. C. Han, Polymer, 2012, 53, 546–554 CrossRef CAS.
  41. B. Almería, W. Deng, T. M. Fahmy and A. Gomez, J. Colloid Interface Sci., 2010, 343, 125–133 CrossRef PubMed.
  42. R. P. A. Hartman, D. J. Brunner, D. M. A. Camelot, J. C. M. Marijnissen and B. Scarlett, J. Aerosol Sci., 2000, 31, 65–95 CrossRef CAS.
  43. S. Freiberg and X. X. Zhu, Int. J. Pharm., 2004, 282, 1–18 CrossRef CAS PubMed.
  44. T.-W. Chung, Y.-Y. Huang and Y.-Z. Liu, Int. J. Pharm., 2001, 212, 161–169 CrossRef CAS PubMed.
  45. M. Louey, M. Van Oort and A. Hickey, Pharm. Res., 2004, 21, 1200–1206 CrossRef CAS.
  46. C. Wang, W. Ye, Y. Zheng, X. Liu and Z. Tong, Int. J. Pharm., 2007, 338, 165–173 CrossRef CAS PubMed.
  47. R. Qi, R. Guo, F. Zheng, H. Liu, J. Yu and X. Shi, Colloids Surf., B, 2013, 110, 148–155 CrossRef CAS PubMed.
  48. F. Zheng, S. Wang, S. Wen, M. Shen, M. Zhu and X. Shi, Biomaterials, 2013, 34, 1402–1412 CrossRef CAS PubMed.
  49. F. Zheng, S. Wang, M. Shen, M. Zhu and X. Shi, Polym. Chem., 2013, 4, 933–941 RSC.
  50. R. Qi, R. Guo, M. Shen, X. Cao, L. Zhang, J. Xu, J. Yu and X. Shi, J. Mater. Chem., 2010, 20, 10622–10629 RSC.
  51. S. Wang, F. Zheng, Y. Huang, Y. Fang, M. Shen, M. Zhu and X. Shi, ACS Appl. Mater. Interfaces, 2012, 4, 6393–6401 CAS.
  52. K. Fu, D. Pack, A. Klibanov and R. Langer, Pharm. Res., 2000, 17, 100–106 CrossRef CAS.
  53. N. Faisant, J. Siepmann and J. P. Benoit, Eur. J. Pharm. Sci., 2002, 15, 355–366 CrossRef CAS PubMed.
  54. Q. Gu, E. Tomaskovic-Crook, R. Lozano, Y. Chen, R. M. Kapsa, Q. Zhou, G. G. Wallace and J. M. Crook, Adv. Healthcare Mater., 2016, 5, 1429–1438 CrossRef CAS PubMed.
  55. S. H. Lim, X. Y. Liu, H. Song, K. J. Yarema and H.-Q. Mao, Biomaterials, 2010, 31, 9031–9039 CrossRef CAS PubMed.
  56. E. Knight and S. Przyborski, J. Anat., 2015, 227, 746–756 CrossRef PubMed.

Footnote

These authors contributed equally to this work.

This journal is © The Royal Society of Chemistry 2017