Nanoparticle-Based Drug Delivery Systems for Neural Interfaces 2020:120

Nanoparticle-based drugNanoparticle-based delivery systems neural for interfaces a novel approach for improved biocompatibility HOLMKVIST ALEXANDER DONTSIOS DEPARTMENT OF EXPERIMENTAL MEDICAL SCIENCE | LUND UNIVERSITY

ALEXANDER DONTSIOS HOLMKVIST Nanoparticle-based drug delivery systems for neural interfaces 2020:120

199831 789176 9

ISSN 1652-8220 ISBN 978-91-7619-983-1 Lund University, Faculty of Medicine Doctoral Dissertation Series 2020:120 Department of Experimental Medical Science Medical Experimental of Department

Nanoparticle-based drug delivery systems for neural interfaces

a novel approach for improved biocompatibility

Alexander Dontsios Holmkvist

DOCTORAL DISSERTATION by due permission of the Faculty of Medicine, Lund University, Sweden. To be defended at Medicon Village lecture hall. 19th of November 2020 at 9.00.

Faculty opponent Professor Jöns Hilborn

3 Organization Document name Doctoral dissertation LUND UNIVERSITY Faculty of Medicine Date of issue Department of Experimental Medical Science 2020-11-19 Neuronano Research Center Author Alexander Dontsios Holmkvist Sponsoring organization

Title and subtitle Nanoparticle-based drug delivery systems for neural interfaces - a novel approach for improved biocompatibility Abstract The overall purpose of this thesis was to reduce brain tissue responses around implanted microelectrodes using a pharmacological strategy. One of the main aims was to develop and evaluate drug delivery systems that allow local administration of anti-inflammatory pharmaceutics. The drug Minocycline was therefore encapsulated into biodegradable poly (D,L-lactic-co-glycolic acid) nanoparticles and thereby also protecting the drug from degradation. The nanoparticle construct resulted in a sustained release of Minocycline over 30 days. This constitutes a substantial increase in release time compared to what has until now been achieved for the drug. The drug-loaded nanoparticles were then embedded in a fast-dissolving gelatin coating surrounding the implant which enabled local and sustained drug release at the target site. This technique supersedes any of the conventional administration routes and minimizes the risk for systemic side effects. The developed drug delivery system was found to significantly attenuate the acute brain tissue responses around implanted microelectrodes in mice. Coatings with Minocycline-loaded nanoparticles significantly reduced the activation of microglia cells compared to control coatings with gelatin alone both 3 and 7 days post implantation. This without affecting the overall microglia population. A significant reduction of the astrocytic response was also found 7 days post implantation in comparison to control implants. No effect on neurons or total cell count was found which may suggest that the Minocycline-loaded nanoparticles are non-toxic to the central nervous system. The thesis also presents a novel nanoparticle-eluting compartmentalized microelectrode that transforms into a flexible tube once implanted.

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Supplementary bibliographical information Language: English

ISSN and key title: 1652-8220 ISBN: 978-91-7619-983-1

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Nanoparticle-based drug delivery systems for neural interfaces

a novel approach for improved biocompatibility

Alexander Dontsios Holmkvist

Cover photo by Alexander Dontsios Holmkvist 2020

Copyright pp 1-55 Alexander Dontsios Holmkvist 2020 Paper 1 © Open access Paper 2 © Open access Paper 3 © by the Authors (Manuscript unpublished)

Faculty of Medicine Department of Experimental Medical Science

ISBN 978-91-7619-983-1 ISSN 1652-8220

Printed in Sweden by Media-Tryck, Lund University Lund 2020

“It´s not who I am underneath, but what I do that defines me”

-Batman

7 Table of Contents

Abbreviations ...... 9 Populärvetenskaplig sammanfattning ...... 10 Papers included in this thesis ...... 12 Introduction ...... 13 Neural interfaces – applications and limitations ...... 13 Foreign body response ...... 13 Minimizing the tissue response ...... 16 Local drug delivery systems ...... 17 Aims ...... 19 Methods and method development ...... 21 Preparation of nanoparticles (I, II, III) ...... 21 Gelatin coating and nanoparticle embedding (II) ...... 25 Surgery and implantation (II) ...... 27 Histology (II) ...... 27 Development of the tubular microelectrode (III) ...... 28 Results and comments ...... 31 1) Minocycline loaded PLGA nanoparticles, in vitro (I) ...... 31 2) Gelatin coating and nanoparticle embedding (II) ...... 34 3) In vivo effects on brain tissue responses (II) ...... 35 4) Development of a tubular microelectrode (III) ...... 38 Discussion ...... 43 In vitro versus in vivo ...... 43 In vivo effects of Minocycline loaded nanoparticles ...... 45 Acute response versus chronic response ...... 46 Concluding remarks and future perspective ...... 47 Acknowledgments ...... 49 References ...... 51

8 Abbreviations

AOT Bis(2-ethylhexyl) sulfosuccinate BBB Blood-brain barrier CNS Central Nervous System DAPI 4′,6-diamidino-2-phenylindole dilactate DLS Dynamic Light Scattering DMAB Didodecyldimethylammonium bromide ESD Emulsification-solvent diffusion FDA Food and drug administration GFAP Glial fibrillary acidic protein GFP Green fluorescent protein HIP Hydrophobic ion pair HPLC High-performance liquid chromatography IL Interleukin MC-NPs Minocycline-loaded nanoparticles NHS N-hydroxysuccinimide ester NeuN Neuronal nuclei NMR Nuclear magnetic resonance Pdi Polydispersity index PFA Paraformaldehyde

PLGA Poly(D,L-lactic-co-glycolic acid) PVA Polyvinyl alcohol ROI Region of interest ROS Reactive oxygen species SEM Scanning electron microscopy TNFα Tumor necrosis factor alpha UV Ultra-violet

9 Populärvetenskaplig sammanfattning

Hjärnan är vårt mest komplicerade organ och det svåraste att studera. Med hjälp av elektroder inopererade i hjärnan är det möjligt att elektriskt stimulera eller registrera in hjärnans elektriska aktivitet. Hjärnimplantaten ger oss därigenom möjligheten att tolka vad nervcellerna försöker åstadkomma i form av t.ex. rörelser. Elektrodimplantat har också många tillämpningar inom sjukvården. De kan t.ex. kopplas till en robotarm eller handprotes som på så sätt kan styras av en person som är förlamad. Informationen kan även användas för att förstå vissa sjukdomstillstånd i nervsystemet och hjälpa personer med olika nerv- och hjärnsjukdomar. Implantat som elektriskt stimulerar vissa områden i hjärnan används redan framgångsrikt för att lindra symptom hos patienter som drabbats av Parkinsons sjukdom. Ett problem med permanent implanterade elektroder är dock att de orsakar en inflammation som kan skada hjärnan. Detta leder till en lokal förlust av nervceller och även bildandet av ärrvävnad runt elektroden. Den cellulära inkapslingen är elektriskt isolerande och resulterar med tiden i en förlust av den kommunikation mellan elektroden och nervcellerna man försöker uppnå. Det övergripande syftet i denna avhandling är att minska den skadliga inflammationen i hjärnvävnaden runt elektroderna med hjälp av läkemedel. Minocyklin är ett läkemedel som skulle kunna vara lämpligt att använda i detta syfte. Läkemedlet är ett antibiotikum med antiinflammatoriska egenskaper och har använts i nästan 40 år för behandling av olika bakterieinfektioner som klamydia, svår akne och tandlossning. Mer nyligen har Minocyklin studerats för dess nervcellsskyddande effekter vid akut hjärnskada, stroke och neurodegenerativa sjukdomar. Läkemedlet måste dessvärre ges i höga doser för att få någon effekt i hjärnan. Ett av syftena i den här avhandlingen var därför att utveckla och utvärdera metoder som möjliggör lokal administrering av antiinflammatoriska läkemedel i hjärnan, dvs kunna tillföra läkemedel direkt till ett specifikt område i hjärnan. Därmed kan man undvika att ge läkemedlet via munnen eller direkt i blodet vilket även påverkar resten av kroppen. Minocyklin kapslades därför först in i nanopartiklar som har en diameter på ca 100 nanometer (1 nanometer = 10-9 meter, ett hårstrå är ca 5000 nanometer i diameter) och är byggda av ett biologiskt nedbrytbart material. Partiklarna bäddades sedan in i en snabbupplösande gelatinbeläggning som omgav hjärnimplantatet. Nanopartikelkonstruktionen visade sig ge en långsam frisättning av Minocyklin under 30 dagar och gelatinbeläggning möjliggjorde en lokal frisättning av nanopartiklar med läkemedel runt hjärnimplantatet precis där det behövs. Denna teknik kompletterar de vanliga intagsvägarna för läkemedel och minimerar risken för biverkningar i resten av kroppen. De utvecklade nanopartiklarna med Minocyklin visade sig påtagligt dämpa inflammationen i hjärnvävnaden kring implanterade elektroder i möss, utan tecken på toxisk vävnadseffekt. Avhandlingen

10 presenterar även en ny slags elektrod som förvandlas till ett flexibelt rör efter implantation. Konstruktionen möjliggör frisläppning av nanopartiklar från en liten öppning i röret, precis där elektroden har kontakt med nervcellerna. Sammanfattningsvis öppnar den nya nanopartikelmetoden upp helt nya möjligheter att bemästra den inflammation som vanligen uppstår runt implanterade elektroder. Detta skulle kunna vara en framtida strategi när det finns behov av permanent implantation av hjärnelektroder.

11 Papers included in this thesis

I. Holmkvist AD, Friberg A, Nilsson UJ, Schouenborg J. Hydrophobic ion pairing of a minocycline/Ca2+/AOT complex for preparation of drug-loaded PLGA nanoparticles with improved sustained release. International Journal of Pharmaceutics. 2016; 499 (1–2): 351-357. II. Holmkvist AD, Agorelius J, Forni M, Nilsson UJ, Linsmeier CE, Schouenborg J. Local delivery of minocycline-loaded PLGA nanoparticles from gelatin-coated neural implants attenuates acute brain tissue responses in mice. Journal of Nanobiotechnology. 2020; 18:27. III. Holmkvist AD, Agorelius J, Retlund A, Nilsson UJ, Schouenborg J. An implantable microelectrode enabling combined local release of drug loaded nanoparticles and sensing via a liquid contact. 2020, in manuscript.

12 Introduction

Neural interfaces – applications and limitations

Neuro-electronic interfaces that connect the human nervous system with computers hold great potential for studying neural mechanisms and helping human patients suffering from neurodegenerative diseases. For this reason, there is a great interest in such interfaces [1]. Neuro-electronic interfaces would allow researchers to record and understand how neuronal networks process information and how this is changed by for example learning or diseases. By recording the signalling of neural networks in the brain and interpreting what the cells are trying to achieve in the form of for example movements, the electrodes can be connected to a robot arm or hand prosthesis and thus controlled by a person who is paralyzed. The interfaces may also be used for stimulation-based therapies, improving the quality of life for patients suffering from neurological injuries or disorders such as chronic pain and Parkinson’s disease. However, in order to study long-term changes in information processing and also enabling treatment of neurological disorders, the foreign body response must be controlled. This response can jeopardize the functions of the implanted neuro-electronic interfaces and the physiological conditions in the adjacent tissue.

Foreign body response

The cells most commonly associated with the brain are neurons, but they only make up for less than 25% of the tissue in the brain. The remaining tissue consists of glial cells (astrocytes, microglia and oligodendrocytes) and vascular-related tissue. As an electrode is inserted into the brain, its path cuts capillaries, extracellular matrix, glial and neuronal cell processes. The mechanical trauma initiates an acute inflammatory process promoting wound closure, neuronal protection, Blood-brain barrier (BBB) repair and restriction of central nervous system (CNS) inflammation. Restoration of the damaged tissue is generally completed within a few weeks after acute uncomplicated insults. However, the reaction changes when a foreign body is left in the tissue and it develops into a chronic response instead. Over time, a coating often referred to as a “glial scar”, forms around the implant that protects the intricate brain

13 structure from the foreign body. This glial scar is detrimental for electrode function and may adversely affect the neuronal population. Microglia and astrocytes are assumed to be predominantly involved in the immune response in the CNS [2]. Microglia are resident macrophages of the brain responsible for clearing cellular debris and toxic substances after injury or during regular cell turnover, thereby maintaining normal cellular homeostasis. They normally reside in an inactive, highly branched state surveying the brain tissue protecting against injury and invasion. When activated, they assume a more compact morphology with phagocytic activity. Immediately after implanting an electrode, microglia begin to extend their processes towards the foreign body and damaged tissue. This fast event is demonstrated by Figure 1, an in-house in vivo two-photon microscopy time series of a laser induced-injury in the cortex in transgenic mice GFP with fluorescently labelled microglial cells (B6129P-CX3CR-1 , Jackson laboratories, USA). Within 30 minutes, the processes of the nearby microglial cells reach the damaged site and appear to fuse together.

0 min 10 min 30 min

GFP Figure 1. Time series of a laser induced injury in the cortex in CX3CR-1 mice with fluorescently labeled microglial cells. Imaged using in-house two-photon microscopy.

Activated microglia up-regulate the production of proteolytic enzymes and pro- inflammatory cytokines including tumour necrosis factor alpha (TNFα), interleukin (IL)-1β, and IL-6 [3]. During inflammation, they also release high concentrations of reactive oxygen species (ROS) including superoxide, hydroxyl radicals, and hydrogen peroxide, all with potentially neurotoxic effects [4]. The released cytokines activate and recruit more microglial cells but also astrocytes. Astrocytes are normally responsible for supplying nerve cells with trophic support, facilitating synapse formation and function, and maintaining the composition of the extracellular fluid of the brain. Upon microglia-induced activation of astrocytes, they may lose many of their normal functions, including the ability to promote neuronal survival and tissue repair. It has also been suggested that activated astrocytes secrete a soluble toxin that rapidly kills neurons and mature oligodendrocytes, but not other CNS cell types [5].

14 The unusual demands of a large injury, such as implanting an electrode, may overwhelm the capacity of microglia and astrocytes. Over the following days and weeks after implantation, microglia and astrocytes migrate to the surface of the implant in an attempt to degrade the foreign object. As the electrodes are constructed to last for a long time and does not simply degrade, it causes a chronic inflammation. Instead of restoring the damaged tissue, these now reactive phenotypes of glial cells exacerbate the inflammatory response and create an unnatural environment for the neurons to function normally. Moreover, aggregated microglia and astrocytes at the surface of the implant may form an electrically insulating sheath that causes a slow dislocation of neurons away from the electrode (Fig. 2). The strength and quality of the recorded electrical signal is highly dependent on the location of the neuron relative to the electrode recording site. The maximum distance action potentials from individual neurons can be recorded at rarely exceeds 50 μm [6]. Glial encapsulation together with altered neurochemical environment will result in poor signal quality and neuronal dysfunction, both detrimental for electrode function.

Acute injury Chronic response

Figure 2. Illustration of the foreign body responses to implantable devices in the brain. The acute injury activate microglia and astrocytes that migrate toward the surface of the device. The chronic response may lead to the formation of a “glial scar” around the device over time. Reprinted (adapted) with permission from (Wellman SM, Kozai TDY. Understanding the Inflammatory Tissue Reaction to Brain Implants to Improve Neurochemical Sensing Performance. ACS Chemical Neuroscience. 2017;8(12):2578-82.) [7]. Copyright (2017) American Chemical Society.

15 Minimizing the tissue response

Electrode design Several studies have demonstrated that electrode design and implantation technique are important factors effecting the tissue response. The brain is in constant movement relative to the skull due to arterial pulsations, respiration, and forces caused by acceleration and rotation of the body. The resulting micro-forces between anchored rigid electrodes and the brain tissue can therefore be expected to stimulate and prolong the activation of glial cells [8]. To make electrodes more biocompatible, highly flexible implants that can follow tissue movements is needed. Implantation of thin flexible electrodes requires some form of structural support to penetrate the brain without bending. Various implantation methods have therefore been developed, for example attaching the implant to a stiffer guide pin that is removed after implantation [9] or adding a hard coating that later dissolves in the brain tissue [10]. Even though these measures have shown to reduce tissue responses, the glial cell responses closest to the electrodes remain to some extent. Using anti- inflammatory pharmaceutics could be a complementary approach to further reduce the tissue reactions.

Minocycline Antibiotic and anti-inflammatory pharmaceutics have widely been used for implantable medical devices such as prosthetic joints, pacemakers and stents. This in order to prevent implant- and surgical site infections [11, 12]. This approach could also be used for neuro-electronic devices as a complementary measure to further reduce the tissue reactions. Minocycline is a widely used broad-spectrum antibiotic that has been in therapeutic use for almost 40 years. It belongs to the second-generation tetracyclines with improved tissue penetration, prolonged half-time and better intestinal absorption in comparison to first generation tetracyclines. The drug has most commonly been used for treatment of bacterial infections such as chlamydia, severe acne, and periodontitis [13, 14]. More recently, Minocycline has been studied for its neuroprotective effects in different animal models of acute traumatic brain injury, stroke and neurodegenerative diseases [15, 16]. It has even shown to be a promising candidate for improving the outcome of neural recordings obtained from a chronic neural interface [17]. The effects are coupled to Minocycline´s ability to inhibit activation and proliferation of microglia [18]. However, to achieve therapeutic drug concentrations in the brain by conventional administration routes, unacceptably large systemic dosages may be needed [19] which could have adverse systemic

16 effects [20]. To improve the therapeutic efficacy and decrease the drug dose required, local delivery from biodegradable nanoparticles could be used. In addition to Minocycline ability to inhibit glial responses, the drug has been in therapeutic use for many years because of its antibiotic properties. Patients who undergo neurosurgery also risk bacterial infections that can cause post- neurosurgical meningitis [21]. Minocycline can therefore play a dual role, both as an anti-inflammatory agent and as an antibiotic prophylaxis, when used in this context with implantation of electrodes in the brain.

Local drug delivery systems

Approximately 98% of small molecular weight drugs and almost 100% of larger molecular weight peptides and proteins do not cross the BBB. In order to deliver therapeutics to the brain more efficiently, nanotechnology-mediated drug delivery systems are emerging. The term “nanoparticle” is a collective name for a particle of any shape with dimensions between 1 and 100 nm [22]. Drug loaded nanocarriers coated with specific compounds can significantly enhance the BBB penetration. A small enough particle coated with for example polysorbate 80 enables passive diffusion over the barrier. If the particles are coated with ligands that bind to proteins associated with the BBB instead, an active transport over the barrier is achieved and may also selectively target for example brain tumor cells [23, 24]. To eliminate the need for repeated administration, the carrier can be made of a degradable material to give a sustained release over a period of days or even weeks after administration. Poly(D,L-lactic-co-glycolic acid) (PLGA) is the most widely used biodegradable polymer in the CNS and undergoes hydrolysis in the body to produce the metabolite monomers, lactic acid and glycolic acid. Since the body effectively deals with these two monomers, there is minimal systemic toxicity associated by using PLGA for drug delivery or biomaterial applications [25]. Even though a nanoparticle strategy for improving BBB permeability offer more effective and non-invasive ways of delivering pharmaceutics to the brain, there is still a lack of specificity to target the area surrounding an implanted electrode. A more straight-forward strategy in this sense would be to deliver the drug-loaded nanoparticles directly from the implant. Delivering drug-loaded nanoparticles locally from implants has been achieved by using different nanoparticle- and coating-materials. Kim et al. [26] for example, used an alginate hydrogel to deliver PLGA particles loaded with the anti- inflammatory agent Dexamethasone. Even though their study showed no significant change in impedance with time in vivo, alginate hydrogels are non-biodegradable in mammalian brains [27]. The coating may thereby create a long-lasting barrier between the implanted electrode and surrounding tissue, hindering neurons getting close enough to the implant for high quality recordings. Mercanzini et al. [28] used

17 a poly(ethylene oxide) coating with embedded poly(propylene sulfide) Dexamethasone-eluting nanoparticles. Their study showed reduced tissue reaction to the implanted microelectrode and that the coating dissolves after implantation. However, the method used for embedding the nanoparticles into the coating involved harsh processing parameters, such as high temperatures, which may be harmful for sensitive molecules. Minocycline for example, degrades easily under exposure to light, high temperatures or low and high pH [29]. In this thesis, gelatin is evaluated as the nanoparticle-carrying coating material. Gelatin is a protein material derived from natural collagen that easily breaks down into amino acids in the body. The material itself has shown beneficial effects for restoration of the BBB after acute brain injury [30] and enables mechanical support for implanting highly flexible electrodes [31]. Therefore, a nanoparticle-based drug delivery system for neural implants should be designed with regard to the following key aspects, i) the drug-loaded nanoparticles should give a sustained release of the drug that matches the inflammatory response, ii) the particles should be large enough to not diffuse freely in the brain but remain close to the implantation site where the highest concentration of the drug is needed, iii) the implant coating should preferably not affect the device function and iv) processing parameters during preparation of both coating and particles needs to be mild enough to protect the drug from degradation.

18 Aims

The overall purpose of this thesis was to reduce brain tissue responses around implanted microelectrodes using a pharmacological strategy. The main aim was to develop and evaluate nanoparticle-based drug delivery systems, suitable for implantable neural devices, that allow local administration of anti-inflammatory pharmaceutics. The specific goals were: 1) To develop and characterize Minocycline loaded PLGA nanoparticles, aiming for particle sizes around 100 nm with sustained drug release over 30 days in vitro. 2) To develop a method for embedding minocycline-loaded PLGA nanoparticles into gelatin-coatings on neural implants. 3) To clarify in vivo whether the addition of these nanoparticles could reduce the acute brain tissue responses after implantation in mice. 4) To develop and characterize a novel nanoparticle-eluting microelectrode in vitro.

20 Methods and method development

This thesis has been a truly interdisciplinary research project that spans organic synthesis, analytical chemistry, surface and colloidal chemistry, drug formulation, neuro-electronic interfaces, small animal surgery, cell biology, immuno- fluorescence staining methods and foreign body responses in the brain. A general description of the methods and techniques used followed by a short discussion of key aspects and motivation within each area is given below. For detailed descriptions of the methods, the reader is referred to the articles this thesis is based on. The roman numbering refers to the different articles included in this thesis.

Preparation of nanoparticles (I, II, III)

Nanoparticles were prepared using a single oil-in-water (o/w) emulsification method often referred to as the emulsification-solvent diffusion (ESD) technique, schematically described in Figure 3. In brief, an o/w emulsion is first made from a partially water-soluble solvent containing both the biodegradable carrier-material and drug, and an aqueous solution containing a surfactant. Water is subsequently added to the two-phase system which causes diffusion of the solvent into the external water phase. This causes the polymer to precipitate in the form of nanospheres with the drug entrapped throughout the particle matrix. The organic solvent is finally evaporated resulting in a colloidal suspension stabilized by the surfactant. All prepared suspensions were lyophilized and stored in a freezer until used. Mannitol and Pluronic F-127 were added as cryoprotectants before freeze- drying to enable redispersion and prevent particle aggregation.

High energy Addition of Solvent emulsification water evaporation Drug+Polymer Organic phase Surfactant Water phase (O/W emulsion) (Solvent diffusion) (Nanoparticles)

Figure 3. Schematic representation of PLGA nanoparticle preparation by the emulsificaton-solvent diffusion technique. Reprinted (adapted) with permission from Paper I.

21 The choice of using the ESD technique for the preparation of nanoparticles in this thesis was based on the method’s compatibility with a wide range of pharmaceutics and its simplicity compared to other existing methods. The method does not require advanced equipment, such as high-pressure homogenizers [32], and can be done with standard laboratory equipment such as an ultrasonic water bath and a magnetic stirrer. It is also known to provide high reproducibility. The resulting nanosphere composition is a homogenous particle with the drug entrapped throughout the particle matrix. This type of particle generally generate longer and more stable release times compared to liposomes or nanocapsules (a solid shell structure that surrounds a core-forming space) prepared by double emulsion techniques [33].

Poly(D,L-lactic-co-glycolic acid) (PLGA), an FDA-approved biodegradable polymer, was chosen as the particulate carrier-material due to its history of safe use in the CNS [34]. PLGA degrades by hydrolysis of its ester linkages in aqueous environments to lactic and glycolic acid which are finally eliminated from the body as carbon dioxide and water. As the polymer degrades and erodes, the drug release kinetics is a function of polymer degradation as well as drug diffusion through polymer matrix. PLGA also enables the possibility to tune the release rate by choosing the composition or ratio between the lactic and glycolic acid monomer. Glycolic acid is slightly more hydrophilic than lactic acid, leading to increased hydrolysis rates and therefore faster release rates with increased glycolic acid content. Didodecyldimethylammonium bromide (DMAB), a double-chained cationic quaternary ammonium surfactant, was chosen as the stabilizing agent due to reports of generating smaller sized nanoparticles compared the more commonly used stabilizer polyvinyl alcohol (PVA) [35]. DMAB also generate particles with cationic surface charge which could increase the rate and extent of their internalization caused from ionic interactions with the negatively charged cell membrane [25]. Different DMAB concentrations were evaluated as a means of adjusting the particle size, aiming for 100 nm.

Hydrophobic ion pairing (I, II) Encapsulating pharmaceutics by the ESD method requires that the drug stays in the oil phase during the preparation and is thus more suitable for lipophilic substances. The initial project plan included screening of suitable immunomodulating drugs, of which many are lipophilic [36, 37], this is the reason why the ESD method was selected. Minocycline is often referred to in the literature as the most lipophilic drugs within the tetracycline family [38], but from a drug formulating perspective it is highly water soluble. In order to encapsulate Minocycline by the ESD technique, the solubility of the drug was altered by the concepts of hydrophobic ion pairing. This involves replacement of the polar counter ions with an ionic detergent. Furthermore, Minocycline is known to be unstable and degrades under e.g. exposure

22 to light, which could lead to degradation of the drug during the preparation process. Advantageously, Minocycline can be stabilized and interacts strongly with calcium ions (Ca2+) [39]. Therefore, a hydrophobic ion pair complex of minocycline, Ca2+ and the anionic surfactant bis(2-ethylhexyl) sulfosuccinate (AOT) was developed (Fig. 4).

Figure 4. Hydrophobic ion pair complex of minocycline, Ca2+ ions and the ionic detergent bis(2-ethylhexyl) sulfosuccinate AOT. Reprinted (adapted) with permission from Paper I.

The formation of the hydrophobic ion pair complex was verified using nuclear magnetic resonance spectroscopy (NMR). The Minocycline/Ca2+/AOT-complex and control samples of Minocycline, Minocycline/Ca2+, AOT, and Minocycline/AOT were dissolved in deuterated chloroform. 1H NMR spectra of each individual sample was recorded on a Bruker Avance II 400MHz spectrometer. The recorded spectra strongly suggested that a Minocycline/Ca2+/AOT-complex was formed and that no interaction occurs between Minocycline and AOT in the absence of Ca2+ ions.

Fluorescently labelled nanoparticles (II, III) Fluorescently labelled nanoparticles were prepared in order to visualize the particles during the development of the gelatin coating and the tubular microelectrode. A fluorescent dye was covalently attached to the PLGA chain [40]. The carboxylic group of PLGA was first activated through the formation of N-succinimide ester (NHS) (Scheme 1). After the reaction, the precipitated biproduct dicyclohexylurea (not shown in scheme) was removed by filtration. The activated polymer was then conjugated with Alexa Fluor 568 cadaverine through an amide bond between the NHS-activated end group of the polymer and the amine group on the fluorescent dye (Scheme 2). Drug-free fluorescently labelled nanoparticles were then prepared according to the particle preparation method described above using fluorescently labelled PLGA mixed with non-labelled PLGA in the ratio of 1:3 (w:w).

Scheme 1. N-hydroxysuccinimide (NHS) activation of the carbocylic endgroup of the PLGA-chain through the reaction with dicyclohexylcarbodiimide (DCC) in anhydrous dioxane.

Scheme 2. Conjugation of NHS-activated PLGA with the fluorescent dye Alexa Fluor 568 in dimethyl sulfoxide (DMSO) with excess amounts of triethylamine (TEA).

Characterization of nanoparticles

Particles size and morphology Dynamic light scattering (DLS) was used for measuring the size, polydispersity index (pdi) and zeta potential of the prepared nanoparticles. This method offers fast analysis and requires minimal sample preparation. The method was therefore ideal for comparing different nanoparticle formulations during the developmental work. Particles suspended in liquid undergo Brownian motion, and the method measures

24 the random changes in the intensity of light scattered from a suspension. The fluctuation in scattering intensity is then converted to particle size through the Stokes-Einstein equation [41]. The instrument therefore reports the size of a hypothetical hard sphere that diffuses in the same fashion as the particle being measured. In order to validate DLS as an appropriate measuring method for our particles, the shape of the nanoparticles was investigated by scanning electron microscopy (SEM).

Drug loading and in-vitro release (I) The amount of Minocycline entrapped in the nanoparticles was quantified using high-performance liquid chromatography (HPLC) with ultra-violet (UV) detection. The in vitro drug release was studied in artificial cerebrospinal fluid using a dialysis method and quantified by HPLC-UV.

Gelatin coating and nanoparticle embedding (II)

Stainless steel needles were used as a model for neural implants. In brief, the needles were first insulated with Parylene-C, dipped in warm gelatin and allowed to dry to create the gelatin coating. The coated needles were then immersed in a nanoparticle suspension at room temperature, allowing the gelatin coating to swell and absorb the suspension without dissolving. The thickness of the coating was measured before and after nanoparticle embedding. The release of fluorescently labelled nanoparticles from coated needles were studied in vitro in agarose gel, a medium widely considered to be a viable model of the brain [42]. Development of the dip-coating method included numerous initial studies on the gelatin solution, including evaluation of different concentrations, temperatures and dipping speed. The absorption of nanoparticles was initially developed by using a casted gelatin rod that was immersed in a suspension of fluorescent nanoparticles in order to visualize how far and fast the nanoparticles were absorbed. A small section of the gelatin rod was then cut and placed between two cover glasses and the cross section was imaged using confocal microscopy (Zeiss LSM 510 with an Acroplan 40×/0.8W objective). Two laser lines, 488 nm and 543 nm were used for sequential excitation of gelatin and the fluorescent nanoparticles, respectively. Figure 5 shows the cross section of a dry gelatin rod after absorption of a nanoparticle suspension with the immersion time of two minutes.

25 A BC

Figure 5. Confocal image of fluorescently labelled nanoparticles (A) absorbed in a gelatin rod (B). (C) is the merged image of (A) and (B).

However, when absorbing particles into gelatin on the coated needles using the same immersion time, the gelatin coatings became deformed after drying. This led to the suspicion that the immersion time was too long for the needles with considerably thinner gelatin coating, causing the gelatin to detach (but not dissolve) from the needle and then dry in a deformed conformation. To find the optimal immersion time, the swelling behavior of gelatin when on needles was therefore studied using light microscopy. Figure 6 shows a time series of a gelatin coated needle after addition of an aqueous solution containing cryoprotectants. The swelling study showed that the coating swelled approximately ten times its original size already after 20 seconds and did not swell substantially more until the coating started to detach from the needle after 45 seconds. The immersion time was therefore set to 30 seconds for absorbing of the nanoparticle suspension. The amount of nanoparticles absorbed in the coating was estimated from the following assumptions: i) the nanoparticles diffuse unhindered into the gelatin coating during swelling and ii) the swelled gelatin coating contained the same concentration of nanoparticles as the surrounding suspension the implant was immersed in.

100 μm 0 s 30 s 60 s

Figure 6. Time series for the swelling of a gelatin coating on a needle. The coating has swelled approximately 10 times its original size after 30 seconds and has started to deform and detach from the needle after 60 seconds.

26 Surgery and implantation (II)

Approval for the animal experiments was obtained in advance from the Malmö/Lund Animal Ethics Committee on Animal Experiments (ethical permit M61-13) and all experiments in this work conform to the regulatory standards of this approval. Transgenic mice (both male and female) that express green fluorescent protein GFP (CX3CR-1 ) in brain microglia were used for the in vivo study on brain tissue responses to implanted needles. The mice were anaesthetized with isoflurane, shaved, and given subcutaneous injection of local analgesia. Small craniotomies (1 mm diameter) were carefully drilled midways between bregma and lambda, around 1 mm laterally of the midline. The coated needles were cut to 3 mm and placed inside a glass capillary filled with paraffin oil to avoid water uptake and swelling of gelatin before entering the cortex. Bilateral implantations of a gelatin-coated needle (control) and a gelatin-coated needle with embedded Minocycline-loaded nanoparticles were done in each mouse. The needles were implanted using a hydraulic micropositioner at the speed of 500 µm/s to a depth of 3 mm below cortical surface.

Histology (II)

The brain tissue responses to the implanted needles were evaluated at 3 or 7 days after implantation. In brief, the mice were deeply anaesthetized by an intraperitoneal injection of pentobarbital and transcardially perfused with saline and paraformaldehyde (PFA). The brains were subsequently removed and post-fixated in PFA overnight. Needles were explanted before the brains were horizontally sectioned (16 μm) on a cryostat.

Immunohistochemistry Two sections from a depth of approximately 400–500 μm into the cortex at each implantation site were then stained using the following primary antibodies: rabbit anti-CD68 to identify activated microglia, rabbit anti-NeuN to identify neurons and chicken anti-GFAP to identify astrocytes. All tissue sections were also stained with 4′,6-diamidino-2-phenylindole dilactate (DAPI) to visualize cell nuclei. As already mentioned, the transgenic mice express green fluorescent protein (GFP) for all brain microglia (that is both resting and activated), characterized by high expression of the chemokine receptor CX3CR-1. Secondary antibodies conjugated to Alexa Fluor 647 and Alexa Fluor 594 were therefore used to enable differentiation of four colors in each section.

27 Image acquisition and analysis The stained sections were imaged using fluorescence microscopy. Regions of interest (ROIs) were set to 0–50 μm and 50–100 μm from the border of the GFP implantation site. The response for all microglia (CX3CR-1 ), activated microglia (CD68) and astrocytes (glial fibrillary acidic protein (GFAP)) was quantified by measuring the proportion of stained area within each ROI. The number of neurons (neuronal nuclei (NeuN) (with a DAPI-positive nucleus as an inclusion criteria) and all cell nuclei (DAPI) were manually counted within each ROI. Mann–Whitney test was used to compare the experimental groups within each ROI and time point, p-values < 0.05 were considered significant.

Development of the tubular microelectrode (III)

The nanoparticle eluting microelectrode was developed from an already existing electrode construction developed in our laboratory. The original microelectrode consisted of a gold wire coated with an electrospun layer of glucose and an insulating Parylene-C coating. The probe has a hole opening in the distal part, allowing fluidic connection with the surrounding tissue. The structure provides sufficient mechanical strength during insertion into the brain but transforms into a flexible tube once implanted. To embed the nanoparticles into the compartmentalized construction, four potential designs (illustrated in Fig. 7) were initially considered and evaluated with regard to electrode function and manufacturing process. All designs comprised a gold wire as the core registering lead and an insulating Parylene-C coating as the outer layer. The designs were the following: A) a glucose compartment with embedded nanoparticles, applied by electrospinning of a glucose solution with suspended nanoparticles. B) electrophoretic deposition of the nanoparticles directly onto the gold lead (inspired by stent technology [43] and making use of the nanoparticles’ positive surface charge) and an outer electrospun compartment of glucose. C) a gelatin compartment with embedded nanoparticles. D) an inner compartment of gelatin with embedded nanoparticles and an outer compartment of electrospun glucose. Characterisation of the tubular electrodes was done using SEM imaging, impedance measurements and fluorescent microscopy. The in vitro release of fluorescent nanoparticles was studied in agarose at 37 °C.

28 A B C D

Parylene Gold wire Glucose Glucose Nanoparticles in glucose Nanoparticles Nanoparticles in gelatin Nanoparticles in gelatin

Figure 7. Tubular electrode designs containing nanoaparticles. All designs comprised a gold wire as the core registrating lead and an insulating Parylene-C coating as the outer layer. A) A glucose compartment with embedded nanoparticles. B) Nanoparticles deposited directly of the onto the gold lead with an outer glucose compartment. C) A gelatin compartment with embedded nanoparticles. D) An inner compartment of gelatin with embedded nanoparticles and an outer compartment of glucose.

30 Results and comments

1) Minocycline loaded PLGA nanoparticles, in vitro (I)

The first aim in this thesis was to develop and characterize Minocycline-loaded PLGA nanoparticles with sizes around 100 nm and an in vitro drug release over at least 30 days. The size of the particles was chosen based on the idea that the particles would remain relatively stationary and not diffuse freely in the brain extracellular space once deposited. The ‘‘pore size’’ of the extracellular space is considered to be ~40 nm [44, 45]. The released drug would thereby be concentrated to the region near the implants where it is most needed. The drug release duration was chosen to match the progression of the brain tissue responses after injury [46]. It has been suggested that Minocycline should be administered at high doses initially for 7 days to attenuate the acute inflammatory response, followed by a lower dose for 3-6 weeks to attenuate the chronic response [47]. Development of the particles started with preparation of drug-free PLGA nanoparticles using the emulsification-solvent diffusion (ESD) technique. Formulations with different concentrations of the surfactant DMAB were evaluated and finally set to 0.10% which resulted in 150 nm sized monodisperse nanoparticles with a positive zeta potential. Lower concentrations resulted in unsuccessful emulsions or larger particles with broader size distributions and negative zeta potentials. Figure 8A shows the size distribution curves generated from dynamic light scattering (DLS) measurements of blank and drug-loaded nanoparticles. The DLS method measures the random intensity changes of light scattered from a suspension. The fluctuation in scattering intensity is then used to calculate the size, poly dispersity index (pdi) and zeta potential of particles within the sample. The zeta potential gives information on how well the particles tend to repel each other and thereby withstand aggregation. Particles with zeta potentials over +30 mV or less than -30 mV will form stable colloidal suspensions. The pdi is a dimensionless measure of the broadness of the size distribution and ranges between 0 and 1. A dispersion may be considered monodisperse (pdi < 0.1) if 90% of the distribution lies within ±5% of the average size.

31 A B

200 nm 200 nm Blank Drug-loaded

Figure 8. A) DLS size distribution curves and SEM micrographs of blank and Minocycline-loaded PLGA nanoparticles. B) In vitro release in artificial cerebrospinal fluid from minocycline loaded PLGA nanoparticles and control curves of minocycline only and minocycline together with cryoprotectants. Reprinted (adapted) with permission from Paper I.

The DLS method reports the size of a hypothetical hard sphere that diffuses in the same fashion as that of the particle being measured. In order to validate the size results obtained with DLS, the shape of the nanoparticles was also investigated by SEM imaging. The SEM micrographs (Fig. 8A) confirm that both the drug-free and drug-loaded nanoparticles were spherical and therefore it was concluded that DLS was an appropriate method to use for characterizing the prepared nanoparticle suspensions. The second step was to encapsulate Minocycline using the ESD technique, a preparation method generally more suitable for encapsulation of lipophilic substances. A novel hydrophobic ion pair (HIP) of Minocycline, Ca2+ and the anionic surfactant bis(2-ethylhexyl) sulfosuccinate (AOT) was developed which enabled encapsulation of the hydrophilic drug using the ESD technique. The formation of the HIP-complex was verified using nuclear magnetic resonance spectroscopy (NMR). The 1H NMR spectra of the HIP-complex and control samples are shown in Figure 9. Preparation of nanoparticles with the encapsulated drug- complex resulted in particle sizes around 220 nm, a drug content of 1.12%, and an entrapment efficiency of 43%. The formulation showed a sustained drug release in vitro over 30 days in artificial cerebrospinal fluid (Fig. 8B).

Figure 9. 1H NMR spectra of (A) Minocycline (MC); (B) MC/Ca2+; (C) AOT; (D) MC/Ca2+/AOT; (E) MC/AOT. Comparison of spectra (A) and (B) show changes in intensity and shape for minocycline’s characteristic peaks at 5.95, 6.83, 7.31, 9.50, 11.53 and 14.87 ppm, which indicates minocycline complexation with Ca2+ ions. Comparison of spectra (C) and (D) show a broadening of AOT’s characteristic peaks at 3.74– 3.98 ppm in the presence of the Minocycline/Ca2+ chelate, strongly suggesting that a Minocycline/Ca2+/AOT-complex has formed. Spectrum (E) of AOT and Minocycline in the absences of Ca2+ appears as a combination of spectra (A) and (C), indicating no interaction occurs between Minocycline and AOT in the absence of Ca2+. Reprinted (adapted) with permission from Paper I.

33 2) Gelatin coating and nanoparticle embedding (II)

The second aim in this thesis was to develop a method for embedding the nanoparticles into gelatin-coatings on neural implants. In order to visualize the particles during the development work (Paper II and III), a fluorescent dye (Alexa Fluor 568) was covalently attached to the PLGA chain before preparing nanoparticles. DLS measurements showed that the fluorescently labelled nanoparticles had similar size, polydispersity index (pdi) and zeta potential as of the nanoparticles prepared in Paper I. The different nanoparticle formulations and characteristics are summarized in Table 1. The similar characteristics between the formulations suggest that the fluorescent dye did not interfere with the formation of nanoparticles and that the fluorescent nanoparticles were a useful surrogate for drug- loaded nanoparticles during the development work.

Table 1 Summary of Size, PDI and zeta potential of the prepared nanoparticles used in this thesis (mean ± s.d., n = 3).

Nanoparticle formulation Size (d. nm) PDI Zeta pot. (mV) Drug free 150 ± 10 0.06 ± 0.020 +57 ±7 Minocycline-loaded 220 ± 6 0.07 ± 0.004 +55 ± 4 Fluorescently labelled 150 ± 2 0.07 ± 0.010 +62 ± 3

An absorption method was developed to embed the nanoparticles into the gelatin coating. Mixing the particles directly into gelatin would have been easier, but as the gelatin solution needs to be heated at 50 °C during the dip-coating step, there would be a risk of melting the particles, degrading the drug and/or causing the drug to be released. The absorption method developed permitted the nanoparticle suspensions to be embedded into the gelatin coating at room temperature. The method also allowed minimal consumption of the laboriously prepared nanoparticle suspension. The final gelatin coatings with embedded nanoparticles had a dry thickness of 9.1 ± 1.2 μm (mean ± standard deviation, n = 7). This thickness was almost doubled compared to the initial gelatin coating alone (4.8 ± 0.9 μm). The amounts of nanoparticles absorbed in the coating was estimated to a total amount of 1 μg nanoparticles on a 3 mm part of the implant. This would correspond to a Minocycline content of 34 ng. Figure 10A shows a needle with fluorescently labelled nanoparticles embedded in the gelatin coating. Figure 10B shows a time series of the in vitro release study in agarose gel and it was concluded that the gelatin coating would stay intact during implantation and that the nanoparticles spread radially ~850 μm from the implant over 10 minutes.

34 A B 4 mm

120 µm 0 min 1 min 10 min

Figure 10. A) A stainless steel needle coated with gelatin and fluorescent nanoparticles. B) Time series of the in vitro release of nanoparticles in agarose gel. Reprinted (adapted) with permission from Paper II.

3) In vivo effects on brain tissue responses (II)

The third aim of this thesis was to clarify in vivo whether the addition of Minocycline-loaded nanoparticles could reduce the acute brain tissue responses after implantation in mice. A study with bilaterally implanted free-floating needles was used as a model for brain implants. The implanted needles had either been coated with gelatin only (control) or gelatin and embedded Minocycline-loaded nanoparticles (MC-NP). The tissue response was quantified in an inner (0–50 μm from border of implant) and an outer (50–100 μm) region of interest (ROI) surrounding the implantation site. Figure 11 and 12 shows representative immunofluorescent images of the brain tissue responses around implantation sites at 3 and 7 days after implantation.

35 Reduced microglia activation Coatings with MC-NPs was shown to significantly reduce the activation of microglia cells (CD68) both 3 and 7 days after implantation compared to the control implants (Fig. 11A-D). After 3 days, the reduction was significant in both inner (p = 0.0079) and outer (p = 0.0052) ROIs. The response had almost disappeared in both groups after 7 days (Fig. 11C, D), but was still significant around implants with MC-NPs compared to the control implants in the inner ROI (p = 0.0289). The overall GFP microglia population (CX3CR-1 ) (Fig 11E-H), however, showed no significant difference between the groups at either time point. These observations suggest that the MC-NPs selectively attenuates the activation of microglial cells without effecting the overall population of microglia.

3 days post implantation 7 days post implantation Control MC-NPs Control MC-NPs CD68 A B C D

GFP E F G H

DAPI I J K L

Merge M N O P

Figure 11. Representative immunofluorescent images of the microglial response. Tissue surrounding gelatin- coated needles with or without embedded MC-NPs at 3 days (two left columns) and 7 days (two right columns) post GFP implantation. Images show activated microglia (CD68) (A–D), CX3CR-1 positive microglia (GFP) (E–H), cell nuclei (DAPI) (I–L), and merge (M–P). Scale bar = 100 μm. Reprinted (adapted) with permission from Paper II.

36 Reduced astrocytic response Both control and MC-NPs implants showed an overall increase in the astrocytic response after 7 days compared to 3 days (Fig. 12A-D). Implants with MC-NPs however, was shown to significantly reduce the astrocytic response in the inner ROI (p = 0.0401) at 7 days compared to the control. It remains to be determined, if this delayed reduction is a direct effect from Minocycline or a secondary effect from the decreased microglia activation at earlier time points.

3 days post implantation 7 days post implantation Control MC-NPs Control MC-NPs GFAP A B C D

NeuN E F G H

DAPI I J K L

Merge M N O P

Figure 12. Representative immunofluorescent images of the astrocytic and neuronal response. Tissue surrounding gelatin-coated needles with or without embedded MC-NPs at 3 days (two left columns) and 7 days (two right columns) post implantation. Images show astrocytes (GFAP) (A–D), neurons (NeuN) (E–H), cell nuclei (DAPI) (I– L), and merge (M–P). Scale bar = 100 μm. Reprinted (adapted) with permission from Paper II.

Neuroprotection Reduction of both activated microglia and reactive astrocytes should be beneficial for neuronal survival as these are known to release components with neurotoxic effects. In this study, we found no difference in neuronal density between the groups at either time point (Fig. 12E-H). This may be due to the gelatin coating itself [48,

37 49] and the study design using free floating implants [50], both of which are beneficial for neuronal survival. Furthermore, there was no significant differences found between groups for all cell nuclei at either time point (Fig. 11I-L and 12I–L). Altogether, the combined findings point to the conclusion that neither the number of neurons, overall microglia population nor the total number of cells are changed by the nanoparticles during the 7-day period. This may suggest that the MC-NPs are non-toxic.

4) Development of a tubular microelectrode (III)

The last aim in this thesis was to develop and characterize a novel nanoparticle- eluting microelectrode in vitro. The four hypothetical designs presented in Figure 7 were evaluated with regard to electrode function and manufacturing process. Design D was found to be the only construction that met the criteria of not affecting electrode function and had a mild enough manufacturing process that did not compromise the integrity of the embedded nanoparticles or drug load. A short comment on each design is given below. Design A was never evaluated due to a risk of compromising the integrity of the nanoparticles and drug during the electrospinning process. The process would use a glucose solution in ethanol and water with suspended nanoparticles which then would be electrospun onto the gold wire. The electrospinning method developed in our laboratory requires high temperature (45-50 °C) for 40 minutes plus additional preheating time. PLGA with a glass transition temperature of 40-45 °C [51] would most likely be affected by the high temperature and might cause the nanoparticles to aggregate and/or release their drug load into the solution [33]. Hence, the strategy of using nanoparticles as a means of sustained release would be lost/compromised and it would also risk that Minocycline would degrade as it is known to be even less stable in solution than a dry powder [39, 52]. Design B was prepared by dragging a gold wire (attached on a frame) through a drop of nanoparticle suspension. Direct current was applied over the syringe needle (cathode) and the wire (anode) and then characterized using fluorescence microscopy. Control experiment using the same setup without current were also made. Preliminary results showed that the nanoparticles were deposited in greater extent when direct current was applied. The full development of this technique was however not completed as characterization or estimation of the deposited amounts of nanoparticles and drug would be challenging and time consuming. Design C was prepared by dip coating the gold wire in gelatin, insulating it with Parylene-C and making a hole opening. Studies on the coating’s integrity was done by immersing the tubes in saline. SEM images (Fig. 13A, B) revealed that the

38 insulation and hole opening of these tubes were compromised due to forces created by the swelling of the gelatin compartment when it absorbs water.

Figure 13. SEM images of tubular microelectrode after immersed in saline. (A) Shank and (B) hole opening of the tubular design with only gelatin as inner compartment. (C) Shank and (D) hole opening of the tubular design with an inner gelatin- and an outer glucose compartment. Modified from Paper III.

Design D was prepared by dip coating the gold wire in gelatin, followed by electrospinning of a glucose layer before insulating with Parylene-C. The design passed the initial integrity test in saline (Fig. 13C, D). It was hypothesised that the quickly dissolving glucose compartment created a space in which gelatin could swell which prevented rupture of the Parylene-C coating. The nanoparticles were embedded into the gelatin layer by absorption as in Paper II. As the wires were held by a frame, the whole frame would need to be immersed into the nanoparticle suspension. This would require a large volume (~30 mL) of suspended nanoparticles and a lot of material would be wasted by absorption into the gelatin coating on the frame´s edges. In order to minimize the consumption of the laboriously prepared nanoparticles, a new method was developed. The wires were instead dragged through a single drop (10 μL) of the nanoparticle suspension. After drying, the outer compartment of electro spun glucose was applied before insulating with Parylene- C.

39 Impedance measurement on the tubular electrode design D showed a quick decrease in impedance, stabilizing at 110 ± 20 kOhm (n = 5) at 1 kHz. It was concluded that the gelatin/nanoparticle coating did not compromise the electrodes functionality and that the design is suitable for electrophysiological recording.

Gelatin + Fluorescent-NPs + Glucose a b c Border

d e f 3.5 mm 3.5

100 μm

Figure 14. Brightfield and fluorescent images of a gold wire after coating with gelatin (a, d); absorption of fluorescent nanoparticles (NPs) (b, e); and electrospinning of glucose with one layer of Parylene-C (c, f). The images were taken at the gelatin dip border (a-c) and at 3.5 mm distance from the dip border (d-f). From Paper III.

Fluorescent and brightfield images of each applied layer (Fig. 14) shows that dip coating the gold wires resulted in a drop shaped layer of gelatin (Fig. 14a, d). The average dry drop diameter of was 22.2 ± 2.2 μm (mean ± standard deviation, n = 9) at the dipping boarder and 19.3 ± 1.7 μm at a 3.5 mm distance from the border. Figure 14(b, e) shows that the fluorescent nanoparticles were successfully absorbed into the gelatin coating without deforming the structure. Electrospinning of glucose and insulating with Parylene-C (Fig. 14c, f) resulted in a wavy outer structure. The gelatin/nanoparticle drops seemed to have increased in size, but this was however rather an optical effect from the rounded structure. It was concluded that the inner gelatin/nanoparticle layer had preserved its structure and that the added glucose had created an outer layer on top of the gelatin.

40 The in vitro release of the fluorescent nanoparticles was studied in agarose at 37 °C. The time series in Figure 15 shows that the fluorescent intensity becomes weaker in both directions from the tube opening as the inside of the tube gets wetted with time. It took approximately 1 hour for the tubes to be completely emptied. It was proposed that water from the surrounding enters the tube and quickly dissolves the glucose followed by gelatin swelling and release of nanoparticles inside the tube. As the tube continues to be filled with water, the gelatin coating dissolves and nanoparticles are released into the surrounding tissue. 1 min 2 min 5 min 10 min 30 min

Tube opening

Figure 15. Image series of the in vitro release of fluorescently labeled nanoparticles from the tubular microelectrode in agarose. From Paper III

42 Discussion

The hypothesis in this thesis was that the biocompatibility of implanted neuroelectronic interfaces could be improved by a pharmacological approach. To this end, novel nanoparticle-based drug delivery systems were developed and evaluated in vitro and in vivo. In the first step, the drug Minocycline was encapsulated into biodegradable PLGA nanoparticles and thereby protected from degradation. This resulted in a sustained release for more than 30 days in vitro. In the next step, the drug-loaded nanoparticles were embedded in a fast-dissolving gelatin coating surrounding the implant thereby enabling local and sustained drug release at the target site. The developed drug delivery system was found to significantly attenuate the acute brain tissue responses around implanted microelectrodes in vivo. In the last step, a novel nanoparticle-eluting compartmentalized microelectrode that transforms into a flexible tube once implanted was developed. Overall, the presented work brings a novel complementary strategy for improving biocompatibility of implanted neuro-electronic interfaces.

In vitro versus in vivo

Before evaluating the effects of the developed drug delivery systems in an animal study, the implants and nanoparticles were characterized in vitro. This was done in order to get a prediction on how the implants and drug delivery systems would behave in vivo and because some experiments are basically unfeasible to do in vivo. An in vitro study also enables comparison to work done by others. The drug content and entrapment efficiency of the nanoparticles developed in this work compares to what others have accomplished for Minocycline-loaded polymeric nanoparticles but the in vitro release time constitutes a substantial increase compared to what has been achieved until now [53, 54]. Comparison of in vitro drug-release profiles with others may however be misleading, as no standard testing procedures are used which may affect the release kinetics. The in vitro test conditions in Paper I (in artificial cerebrospinal fluid at physiological pH and temperature) were chosen to mimic the prospective environment the nanoparticles were intended for. However, since the brain is an extremely complex environment it is difficult to cover all aspects.

43 The in vitro study in Paper II (in an agarose gel at room temperature) showed that the gelatin coating could withstand an implantation and also gave an idea of the nanoparticle release from the implant. The choice of doing the study at a lower temperature instead of at physiological temperature was done with regard to suggestions that the brain surface temperature is closer to room temperature during a craniotomy were the insulating layer of hair and skin has been removed [55]. No attempt was made to characterize the drug release in detail. The in vitro study in Paper III was also done in an agarose gel, but in this study at a physiological temperature to mimic the brain environment around an implant after surgery. The study showed the nanoparticle release from the tubular electrode and that the implant could withstand implantation at an elevated temperature. Even though agarose gels are widely considered to be a viable in vitro model of the mechanical strength of the brain [42], it should be noted that diffusion of nanoparticles might differ substantially in the extracellular space compared to that in agarose [56, 57]. The “pore size” is bigger in agarose compared to in the extracellular space and the particles will therefore diffuse more freely in the model compared to in a live situation. We also tried to examine the in vivo spreading of nanoparticles in a pilot study using gelatin coated implants with embedded fluorescent nanoparticles and looking for them 3 days after implantation. However, they were not to be found with either fluorescent or confocal microscopy. It was hypothesized that the amount of nanoparticles became too small in the tissue and the emitted fluorescence was below detection limits. Videos from the in vitro release studies of nanoparticles in agarose in paper II and III supported this notion. In paper II, it was clearly seen that the fluorescence became weaker the further away the particles came from the implant. In paper III, the particles could not even be seen after leaving the gelatin. Another possible “fate” in vivo is that the particles are phagocytosed by nearby microglia in which degradation of the particles might be faster compared to the in vitro milieu which would give a faster drug release kinetic. As evident, there are a number of possibilities and aspects to consider when setting up in vitro studies. It comes to a point where further in vitro tests, evaluations and predictions will not give more insights for proving the hypothesis. A faster and perhaps more valid approach thereafter may be to study the effects in vivo instead. The overall purpose of this thesis was nevertheless to reduce brain tissue responses around implanted microelectrodes using a pharmacological strategy. In Paper II we showed that the developed nanoparticle drug delivery system indeed reduced the acute tissue reactions around implanted microelectrodes. The mechanisms behind the effects will however need further studies.

44 In vivo effects of Minocycline loaded nanoparticles

In Paper II it was shown that the Minocycline-loaded nanoparticles significantly reduced the microglial activation both after 3 and 7 days, and the astrocytic response after 7 days. These findings confirm and extend previously reported effects of Minocycline on microglia [58] and astrocytes [17]. It remains to be determined however, if the delayed effects on astrocytes found is a direct effect from Minocycline or a secondary effect from the decreased microglia activation at earlier time points. It has been suggested that activated microglia polarize into two distinct functional phenotypes, pro-inflammatory (M1) and anti-inflammatory (M2). The M1- phenotype is associated with production of proinflammatory cytokines and chemokines that are involved in cytotoxicity and microbial killing. The M2- phenotype is associated with the expression of scavenger receptors and pro- angiogenic factors that promotes tissue repair and functional recovery [59, 60]. Minocycline’s inhibitory effects on microglial activation has been readily studied and reports have also shown that Minocycline selectively inhibit microglial polarization into the M1 subgroup, thereby increasing the relative importance of the anti-inflammatory subgroup M2 [61, 62]. Evidence for that astrocytes also polarize into at least two types of phenotypes with different properties is emerging [63]. These are termed A1 and A2, one type being harmful and the other helpful for CNS recovery and repair. A1 astrocytes are neurotoxic and upregulate cascade genes shown to be destructive to synapses. A2 astrocytes upregulate neurotrophic factors that are shown be protective after injury. It has further been suggested that activated microglia induce A1 astrocytes [5]. The antibody (CD68) used to detect activated microglia in our study, stain both the M1 and the M2 microglial subgroups, but not unactivated microglia. The reduction of activated microglia is therefore likely related to a reduced number of M1 microglia given the established pharmacological mechanism of Minocycline. As a consequence, the number of activated astrocytes would also be reduced. Given the enormous complexity of the neuronal tissue, while possibly catching some of the mechanisms, this explanation of Minocycline´s effects is likely to be too simplistic. The situation is probably even more complex if other efforts involving changes in configuration and size of the electrodes [48, 64] are undertaken in order to reduce the acute tissue responses.

45 Acute response versus chronic response

Minimizing tissue responses from implanted neuroelectronic devices is necessary for obtaining high quality recordings that are stable for long periods of time. Studies have shown that the mechanical properties such as shape, flexibility and anchoring of the electrode are of great importance for lessening the chronic responses [1]. Moreover, the initial injury caused by the surgery and implantation itself has proven to be an important factor on the later chronic response [7]. Kumosa et al. [30] for example demonstrated that the astrocytic response in rat cortex can be seen for at least 6 weeks after a stab wound injury in the absence of a chronic implant. The study also showed that a gelatin coating could reduce this response significantly. The recently developed tubular microelectrode that is stiff enough to be implanted and then transforms into a flexible probe has proved to be highly biocompatible and essentially eliminates the loss of nearby neurons (Agorelius et al., 2020 unpublished). However, 6 weeks after implantation some remnants of astrocytic and microglial cells were observed around but also inside the tubes. To further reduce the acute responses using a pharmaceutical approach as presented in this thesis brings in an additional factor. Paper II focused on responses that occurred during the first 7 days after implantation and both free floating implants and a gelatin coating were used in the study. As already mentioned, these two measures have proved to be important factors in reducing both acute and chronic responses [10, 50, 64]. The study demonstrated that Minocycline inhibited microglia activation and reduced the astrocytic response but showed no observable effects on neuronal survival. The lack of observable effects on neuronal survival might be due to that the neuroprotective effects of Minocycline are masked by improvements related to electrode design or that the effects of Minocycline on neuronal survival occur at later time points. The long-term effects of pharmaceutically reducing the acute responses needs to be evaluated further. One could speculate that the situation could be altered in the chronic phase. Minocycline and other anti-inflammatory drugs interfere after all with the CNS´s own protective and repairing mechanisms. Under normal circumstances, it might be harmful to disturb these balances. The brain may lose its own ability to for example fight off infections or remove dysfunctional neurons to preserve neuronal circuit functions. However, it is no longer a normal situation when an electrode is in the brain. Then it may be an idea to pharmaceutically modulate these mechanisms with the aim of maintaining that balance. Another speculation is that to much of M2 and lack of M1 microglia may lead to an excessive reconstruction of the tissue and result in an even thicker scar around the implant. These speculative scenarios would most likely also be dose dependent.

46 Concluding remarks and future perspective

The novel drug-delivery-systems for neural interfaces developed here provide highly localized effects in the tissue at the very electrode contact. By practically eliminating risks for systemic side effects this greatly improves the drug´s therapeutic efficacy as compared to systemic administration. Moreover, the tubular electrode construction developed in Paper III offers controlled release of biologically active compounds into the surrounding tissue directly outside the tubular orifice and recording site. This may be used to pharmacologically preserve the tissue of interest and inhibit migration of glial cells into the electrode to reduce the risk of clogging. While focusing on minocycline in this thesis, it is obvious that a similar approach may also be useful for other pharmaceutics and/or applications for which a highly localised effect is advantageous. The design also opens the possibilities to deliver and analyze highly localized effects of other drugs such as neurotrophic factors, with reduced confounding factors. From a drug development point of view, the hydrophobic ion pairing concept presented in Paper I may also be applicable to other water-soluble drugs and thereby enable use of preparation methods which are otherwise only suitable for lipophilic substances. As a final note, it should be kept in mind that the brains immune system is truly an extraordinary complex system, developed through many years of evolution. Pharmaceutically altering this system should be done with great caution and with the aim of helping the brain maintain its ability to protect and repair itself.

47 48 Acknowledgments

I wish to express my sincere gratitude to: Professor Jens Schouenborg, for introducing me into the field of neurophysiology and for stimulating scientific guidance. Professor Ulf Nilsson, for stimulating collaboration, chemical guidance and support throughout this work. Professor Nils Danielsen, for encouragement and support throughout this work. Associate professor Cecilia Eriksson, for assistance with immunohistochemistry and support.

To all co-authors, former and present colleagues and office roommates, technical staff and administrative personal at the Centre for Analysis and Synthesis (CAS) and at Neuronano Research Center (NRC), for all help and good times at and after work.

To my family and friends, for your patience, love and support.

50 References

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Paper I

International Journal of Pharmaceutics 499 (2016) 351–357

Contents lists available at ScienceDirect

International Journal of Pharmaceutics

journal homepage: www.elsevier.com/locate/ijpharm

Pharmaceutical Nanotechnology

Hydrophobic ion pairing of a minocycline/Ca /AOT complex for

preparation of drug-loaded PLGA nanoparticles with improved

sustained release

a,b, a b a

Alexander Dontsios Holmkvist *, Annika Friberg , Ulf J. Nilsson , Jens Schouenborg

Neuronano Research Center, Department of Experimental Medical Science, Medical Faculty, Lund University, Sweden

Centre for Analysis and Synthesis, Department of Chemistry, Lund University, Sweden

A R T I C L E I N F O A B S T R A C T

Article history: Polymeric nanoparticles is an established and efﬁcient means to achieve controlled release of drugs.

Received 25 September 2015

Incorporation of minocycline, an antibiotic with anti-inﬂammatory and neuroprotective properties, into

Received in revised form 28 December 2015

biodegradable nanoparticles may therefore provide an efﬁcient means to combat foreign body reactions

Accepted 5 January 2016

to implanted electrodes in the brain. However, minocycline is commonly associated with poor

Available online 7 January 2016

encapsulation efﬁciencies and/or fast release rates due to its high solubility in water. Moreover,

minocycline is unstable under conditions of low and high pH, heat and exposure to light, which

Chemical compounds studied in this article:

exacerbate the challenges of encapsulation. In this work drug loaded PLGA nanoparticles were prepared

PLGA (PubChem CID: 23111554)

by a modiﬁed emulsiﬁcation-solvent-diffusion technique and characterized for size, drug encapsulation

DMAB (PubChem CID: 18669)

and in vitro drug release. A novel hydrophobic ion pair complex of minocycline, Ca ions and the anionic

AOT (PubChem CID: 23673837)

Minocycline (PubChem CID: 54685925) surfactant AOT was developed to protect minocycline from degradation and prolong its release. The

optimized formulation resulted in particle sizes around 220 nm with an entrapment efﬁciency of 43% and

Keywords: showed drug release over 30 days in artiﬁcial cerebrospinal ﬂuid. The present results constitute a

Minocycline

substantial increase in release time compared to what has hitherto been achieved for minocycline and

Poly(D,L-lactic-co-glycolic acid) (PLGA)

indicate that such particles might provide useful for sustained drug delivery in the CNS.

Nanoparticles

ã 2016 The Authors. Published by Elsevier B.V. This is an open access article under the CC BY license

Drug release

(http://creativecommons.org/licenses/by/4.0/).

Hydrophobic ion paring

Emulsiﬁcation-solvent-diffusion method

1. Introduction barrier (Pardridge, 2005) and short half-life (Di Stefano et al.,

2008).

Minocycline, a broad-spectrum antibiotic with anti-inﬂamma- To improve the therapeutic efﬁcacy and decrease the drug dose

tory properties, has been shown to have neuroprotective and required from minocycline and other water-soluble drugs, local

neurorestorative effects in ischemic injury/stroke and neurode- delivery from biodegradable nanoparticles could be used (Orive

generative disease models (Lee et al., 2004; Ryu et al., 2004). The et al., 2009). One of the most widely used biodegradable material

mechanisms behind these effects are believed to lay in minocycline for drug delivery is poly(D,L-lactic-co-glycolic acid) (PLGA) as it

’s ability to inhibit proliferation and activation of reactive microglia undergoes hydrolysis in the body to produce the metabolite

(Giuliani et al., 2005; McAllister Ii and Miller, 2010; Tikka et al., monomers, lactic acid and glycolic acid. Since the body effectively

2001). The neuroprotective role of minocycline has been used to deals with these, there is minimal systemic toxicity associated

improve the quality and longevity of chronic neural recordings by with using PLGA for drug delivery even in the CNS (Houchin and

reducing glial sheath growth on implanted microelectrodes in the Topp, 2008; Kumari et al., 2010). The PLGA particles can be

central nervous system (CNS) (Rennaker et al., 2007). However, the designed to give a sustained release over a period of days or even

effects were only seen at high doses by conventional administra- weeks after implantation. Hence, such a sustained release of

tion due to the molecules restricted passage over the blood–brain- minocycline from nanoparticles would be ideal for reducing the

foreign body response seen after implanting electrodes in the brain

(Polikov et al., 2005).

Encapsulation of water-soluble drugs into PLGA nanoparticles is

* Corresponding author at: Neuronano Research Center, Department of

generally done by double emulsion techniques (Astete and Sabliov,

Experimental Medical Science, Medical Faculty, Lund University, Medicon Village,

ﬁ

404 A2, Scheelevägen 2, SE 223 81 Lund, Sweden. 2006). This approach often yields high encapsulation ef ciencies

E-mail address: [email protected] (A.D. Holmkvist).

http://dx.doi.org/10.1016/j.ijpharm.2016.01.011

0378-5173/ã 2016 The Authors. Published by Elsevier B.V. This is an open access article under the CC BY license (http://creativecommons.org/licenses/by/4.0/).

352 A.D. Holmkvist et al. / International Journal of Pharmaceutics 499 (2016) 351–357

but has the drawback of creating porous particles that are rapidly minocycline with Ca signiﬁcantly enhances the stability of the

wetted upon contact with aqueous media, which subsequently drug in solution (Soliman et al., 2010).

results in faster release rates (Fredenberg et al., 2011; Siepmann Unlike other tetracyclines, minocycline contains two amino

et al., 2005). Using homogenous particles where the drug is groups that are responsible for its high solubility in water. In total,

entrapped throughout the particle matrix is preferable in this it has four ionizable groups with the dissociation constants of

sense (Klose et al., 2006; Stevanovic and Uskokovic, 2009). pKa1 = 2.8; pKa2 = 5.0; pKa3 = 7.8; and pKa4 = 9.3 (Kalish and Koujak,

Preparation of these is usually based on single oil-in-water (o/ 2004). The isoelectric point of minocycline is at pH 6.4 were the

w) emulsiﬁcation methods, such as the emulsiﬁcation-solvent- molecules neutral zwitterionic form is predominant and the

diffusion (ESD) technique (Kwon et al., 2001; Pinto Reis et al., 2006; highest partition coefﬁcient in octanol is found (Zbinovsky and

Quintanar-Guerrero et al.,1997) schematically shown in Fig.1. Both Chrekian, 1977). The high aqueous solubility of ionic compounds

drug and polymer are ﬁrst dissolved in an organic solvent followed stems not only from being charged, but also from the fact that the

by emulsiﬁcation into an aqueous phase containing a surfactant. A counter ions easily solvates by water. If the counter ion would be

large amount of water is then added to the system leading to replaced with a species with similar charge, but less solvated by

solvent diffusion to the external phase and the formation of water it is expected that the aqueous solubility would decrease.

nanospheres. Finally the solvent is evaporated, resulting in a stable This process is referred to as hydrophobic ion pairing (HIP)

colloidal suspension containing nanoparticles of solid polymer (Jacobson et al., 2010; Meyer and Manning, 1998) and involves

stabilized by the surfactant. The surfactant type and concentration stoichiometric replacement of polar counter ions with ionic

will have a large impact on the ﬁnal particle size and affects the detergents of similar charge.

drug loading efﬁciency (Gómez-Gaete et al., 2007), sphericity The purpose of this study was to prepare DMAB stabilized

(Esposito et al., 1999) and surface charge of the formed particles minocycline-loaded PLGA nanoparticles by the ESD-method,

(Italia et al., 2007). A commonly used surfactant for this aiming for particle sizes around 100 nm, high encapsulation

preparation method is didodecyldimethylammonium bromide efﬁciencies and long release times. We have evaluated two

(DMAB), a double-chain compound belonging to the class of approaches to increase minocycline’s solubility in organic solvents

cationic quaternary ammonium surfactants (Kwon et al., 2001). by replacing the water-soluble hydrochloride salt form of

Cationic surface charges are desirable as it promotes interaction of minocycline with either its isoelectric form at pH 6.4 or a HIP-

the nanoparticles with cells and hence increases the rate and complex of minocycline, Ca and the anionic surfactant bis(2-

extent of internalization (Kumari et al., 2010). The ESD technique ethylhexyl) sulfosuccinate (AOT), an FDA approved food additive

presents several advantages, such as high batch-to-batch repro- (Fig. 1C). The prepared nanoparticles were characterized for size,

ducibility, simplicity and narrow size distribution. Disadvantages morphology, drug loading and in vitro drug release.

are the large volumes of water to be eliminated from the

suspension and the leakage of water-soluble drugs into the 2. Materials and methods

aqueous external phase during emulsiﬁcation. Drug leakage will

result in low encapsulation and can lead to a high initial burst All samples were prepared in triplicate and tested with triple

release (Danhier et al., 2012) if the leaked drug adsorbs onto the measurements or more. Data is presented as the mean standard

surface of the particles rather than being dispersed throughout the deviation of the mean.

polymer matrix. Another contributing factor for low encapsulation

efﬁciency is degradation of the drug during preparation (Xu et al., 2.1. Materials

2009). Minocycline is known to be unstable under acidic or

alkaline conditions, heat and exposure to light, where it undergoes Poly(D,L-lactic-co-glycolic acid) (PLGA) (50:50), Mw 24000-

epimerization and degrades (Chow et al., 2008; Zbinovsky and 38000 acid terminated (Sigma). Minocycline hydrochloride

Chrekian, 1977). Studies have however shown that chelation of (Sigma–Aldrich). Didodecyldimethylammonium bromide (DMAB)

Fig. 1. (A) Schematic representation of PLGA nanoparticle preparation by the ESD-technique. (B) Proposed composition of the formed drug-loaded nanoparticles; the

entrapped drug is distributed throughout the polymer matrix and the particles surface is covered with the cationic surfactant DMAB. (C) HIP-complex of minocycline, Ca

ions and bis(2-ethylhexyl) sulfosuccinate (AOT).

A.D. Holmkvist et al. / International Journal of Pharmaceutics 499 (2016) 351–357 353

(Sigma–Aldrich) was used as stabilizer. Sodium bis(2-ethylhexyl) use of Tris–HCl buffer (10 M, pH 7.4) instead of bis–tris buffer in

sulfosuccinate (AOT) (D4422, Sigma–Aldrich) was used as an ion all aqueous solutions. Five different D/P ratios (1, 1.5, 2, 3 and 4%)

pairing agent. HEPES-buffered artiﬁcial cerebrospinal ﬂuid pH 7.3 were evaluated.

(aCSF) (Protocols, 2011) was used as drug release medium. All

other chemicals and solvents used were of HPLC grade. Milli Q 2.3. Particle size and morphology

ultra-pure water (Millipore) was exclusively used for the

preparation of all aqueous solutions. The hydrodynamic size, polydispersity index (PDI) and zeta

potential of the prepared nanoparticles was determined by

2.2. Preparation of nanoparticles dynamic light scattering (Malvern Zetasizer Nano-ZS), taking the

average of three measurements for each batch. The suspensions

Nanoparticles were prepared by an ESD technique (Quintanar- were diluted ten times with the aqueous phase used in the

Guerrero et al., 1997). In brief, PLGA (20 mg) was dissolved in an formulation before measurements. The morphology of the nano-

ethyl acetate/methanol mixture (9:1, 1 ml) at room temperature. particles was investigated by SEM (FEI Nova NanoLab 600) at a

The organic phase was added drop-wise to a bis–tris buffered working distance to 5.2 mm and an accelerating voltage of 15 kV.

aqueous solution (10 mM, pH 6.4) containing a stabilizer (DMAB, The sample was prepared by spin-coating the ﬁltered crude

0.10%, 1.2 ml) under magnetic stirring at 600 rmp. The two-phase suspension on a silicon wafer at 2000 rpm for 1 min.

system was then emulsiﬁed by sonication for 10 minutes using an

iced ultra-sonic cleaning bath (VWR USC 300 D, 80 W at 45 kHz). To 2.4. Drug encapsulation

this emulsion, water (12 ml) was added under constant stirring,

which resulted in nanoprecipitation. The suspension was then kept Centrifugation was used to separate the drug-loaded nano-

in an open beaker with magnetic stirring (400–500 rpm) overnight particles from the residual free drug in the suspensions. All

to evaporate the organic solvent. To prepare drug-loaded nano- nanoparticle suspensions were ﬁltered through a sintered glass

particles, minocycline was dissolved together with PLGA in the ﬁlter before the separation. The suspensions were centrifuged

organic phase following the same protocol. The inﬂuence of (Centrifuge 5430, Eppendorf) at 7197 rcf for 30 min, the superna-

altering the following formulation parameters was investigated tant was discarded and the pellet was resuspended in pure water,

and optimized with respect to particle size and drug incorporation, three times. The washed nanoparticles were ﬁnally frozen in liquid

aiming for particle sizes around 100 nm and high encapsulation nitrogen and lyophilized at 55 C and 0.050 mbar for 12 h

efﬁciency. (Freezone 4.5 model 77510, Labconco). The freeze-dried nano-

DMAB concentration particles (NPs) were weighted and dissolved in acetonitrile (0.1%

Blank nanoparticles were prepared using different DMAB TFA, 400 ml). Methanol (0.1% TFA, 1 ml) was then added to dissolve

concentrations (0.01–0.15%) in bis–tris buffered water phase the drug and precipitate the polymer. The solution was centrifuged

(10 mM, pH 6.4) and then evaluated for size and zeta potential. at 1000 rcf for 5 min and the supernatant was ﬁltered through a

Preparation of nanoparticles with the isoelectric form of minocy- 0.45 mm Teflon filter. The amount of minocycline was quantified by

cline at pH 6.4 HPLC (Hewlett-Packard LC 110 0 Series, Agilent Technologies,

Minocycline was dissolved in water and set to its isoelectric Foster City, CA, USA) at 25 C using a Xterra MS 50 2.1 mm,

point at pH 6.4 by addition of 0.2 M NaOH. Minocycline at its 2.5 mm particle size reversed phase C18 column (Waters, Millford,

isoelectric point was extracted with dichloromethane (DCM) and CT, USA). A mobile phase gradient program was used starting at

the solvent was evaporated under reduced pressure. Particle 95% triﬂuoroacetic acid (0.05%), gradually changing over to 30%

preparation was performed as described above and three different acetonitrile in 4 min and then ﬁnishing by changing to 95%

drug-to-polymer (D/P) ratios (5, 12.5 and 20%) were evaluated. acetonitrile in one minute at the ﬂow rate of 0.27 ml/min. An

Preparation of nanoparticles with a HIP complex of minocycline, injection volume of 5 ml and a detection wavelength of 354 nm

2+ 2

Ca and AOT were used. Linear calibration curves (r 0.999) were obtained

The HIP complexation was accomplished according to Soliman over the drug concentration range of 2–20 mg/ml.

et al. (2010) and the Bligh-Dyer method (Bligh and Dyer, 1959) in a The drug encapsulation is expressed both as drug content (DC)

monophase of dichloromethane, water, and methanol. Stock and entrapment efﬁciency (EE), represented by Eqs. (1) and (2) and

solutions of CaCl2 and minocycline were prepared in Tris–HCl respectively.

buffer (10 M, pH 7.4). The pH of the solutions was adjusted to

mass of drug in NPs

7.4 using 0.1 M NaOH, if necessary. The CaCl2 solution was added to DC ð%Þ ¼ 100 ð1Þ

mass of freeze dried NPs

the minocycline solution at speciﬁc volume to attain a minocy-

cline/Ca molar ratio of 1:2. This solution was then added to an

equal volume of DCM containing an equimolar amount of AOT with

2+ actual drug content

ð%Þ ¼ ð Þ

respect to the Ca ions. A mono-phase was formed by addition of EE 100 2

theoretical drug content

methanol at volume ratios of 1:1:2.1 of (buffer/DCM/methanol).

The solution mixture was ﬁnally phase separated by addition of

equal volumes of buffer (1 ml) and DCM (1 ml) followed by 2.5. Drug release

centrifugation at 1000 rcf for 1.5 minutes, with the minocycline/

2+ 2+

Ca /AOT-complex now present in the organic phase. The complex The drug release from the optimized 3% minocycline/Ca /AOT-

was isolated by separation of the two phases and evaporation of formulation was studied in vitro using a dialysis technique. Three

the solvent. H NMR spectra of the formed complex were recorded batches of freshly prepared nanoparticle suspensions (total

on a Bruker Avance II 400 MHz spectrometer in deuterated volume of 42 ml, containing 60 mg PLGA and 1.8 mg minocycline)

ﬁ ﬁ

chloroform (CDCl3) at room temperature. Chemical shifts are were combined, ltered through a glass lter and lyophilized.

given in ppm downﬁeld from the signal for Me4Si, with reference to Mannitol (2% w/v) and Pluronic F-127 (1%, 3 ml) were used as

residual CDCl3. Control samples of minocycline, minocycline/Ca cryoprotectants to enable redispersion and prevent particle

(molar ratio 1:2), AOT, and minocycline/AOT (molar ratio 1:2) all aggregation. The freeze-dried nanoparticles were resuspended

subjected to the same method were also recorded. Particle in 5 ml of aCSF and placed in a dialysis bag (SnakeSkin, 10K MWCO,

ﬁ

preparation was performed as described above except for the Thermo Scienti c, Rockford, IL, USA). The dialysis bag was then

354 A.D. Holmkvist et al. / International Journal of Pharmaceutics 499 (2016) 351–357

immersed in a amber beaker containing 35 ml aCSF and main- Generally, PLGA nanoparticles generate negative zeta potentials

tained at 37 C in a water bath. At predetermined time intervals, the due to the PLGA’s free carboxylic end groups (Kuo and Yu, 2011;

whole release medium was replaced with fresh medium. The Stolnik et al., 1995). Here, the zeta potential shifted to positive

released amount of minocycline was quantiﬁed by HPLC at 25 C values with increasing surfactant concentration most likely as an

using a Sunﬁre 100 3.0 mm, 3.5 mm particle size reversed phase effect of the increasing surface coverage of DMAB on the prepared

C18 column (Waters, Millford, CT, USA). An isocratic mobile phase particles.

of triﬂuoroacetic acid (0.05%)-acetonitrile at the volume ratio of It should be noted that the 0.05% and 0.075% formulations

87:13 and ﬂow rate of 0.6 ml/min was used. All samples were initially formed visibly stable suspensions. These suspensions,

filtered through a 0.20 mm Teflon filter. An injection volume of however, started to flocculate with time and the reported sizes are

20 ml and a detection wavelength of 354 nm were used. Linear more likely of agglomerated nanoparticles, as indicated by their

calibration curves (r 0.980) were obtained over the drug higher PDI-values. A possible explanation is that the amount of

concentration range of 1–10 mg/ml. Evaluation of the dialysis surfactant was sufﬁcient to stabilize the dispersed droplets of

method described above was done by measuring minocycline’s solvent in the o/w emulsion and enable the nanoparticle formation

diffusion over the dialysis bags in two different setups; minocy- during preparation but insufﬁcient to stabilize the formed

cline only and minocycline together with cryoprotectants in aCSF. nanoparticles in the ﬁnal suspension. Therefore the zeta potential

The stability and degradation of minocycline in aCSF (25, 50 and for these formulations more likely represents that of the formed

100 mg/ml) under the in vitro release conditions was also agglomerates and not of the nanoparticles.

investigated in a seven-day study. The surfactant concentration should be kept to a minimum to

reduce the risk of interactions between drug and surfactant, but yet

3. Results and discussion high enough to prevent particle aggregation stabilize the

suspensions. A DMAB concentration of 0.10% was therefore used

3.1. Particle size and morphology in all subsequent formulations. Our optimized HIP-complex

formulation with a theoretical D/P ratio of 3% generated a positive

Initial studies were made on the surfactants effect on size and zeta potential of 55 4 mV and somewhat larger particles as

morphology. Drug free nanoparticles were prepared using differ- compared to the blank formulation (220 nm and 150 nm,

ent DMAB concentrations and evaluated for size, PDI and zeta respectively). This is illustrated by the size distribution curves

potential. The PDI is a dimensionless measure of the broadness of generated via DLS in Fig. 2. The more or less unchanged zeta

the size distribution calculated from a cumulants analysis of the potential of the HIP-complex formulation compared to the blank

DLS-measured intensity autocorrelation function. In the Zetasizer (+55 mV and +57 mV, respectively), indicate that there was no

software the PDI ranges from 0 to 1. Numerically, a dispersion may interaction between drug and particle surface.

be considered monodisperse if 90% of the distribution lies within All size measurements were performed using DLS, which

5% of the average size. The PDI typically has values less than reports the size of a hypothetical hard sphere that diffuses in the

0.1 for a monodisperse test sample (ISO-22412:2008). Values same fashion as that of the particle being measured. In order to

greater than 0.7 indicate that the sample has a very broad size validate the size results obtained with DLS, the overall shape of the

distribution and is probably not suitable for the dynamic light nanoparticles was investigated by SEM imaging. The SEM micro-

scattering (DLS) technique. As can be seen in Table 1, formulations graphs in Fig. 2 conﬁrm that both the blank and the HIP-complex

with DMAB concentrations lower than 0.10% resulted in larger loaded nanoparticles were spherical.

particles with broader size distributions (PDI > 0.1), negative zeta

potentials and/or unsuccessful emulsions. Formulations with

DMAB concentrations of 0.10% and higher resulted in monodis-

perse nanoparticles (PDI < 0.1) with positive zeta potentials. The

mean particle size was 150 nm in diameter and decreasing with

increasing surfactant concentrations.

The amount of surfactant used during particle preparation is an

important factor for both size reduction and stability of the

prepared nanoparticle suspensions. Too low concentration leads to

aggregation of the polymer, whereas too high concentration may

affect drug incorporation as a result of the interaction between

drug and stabilizer. The zeta potential gives information on the

stability of the prepared nanoparticle suspensions and is strongly

inﬂuenced by the particles surface charge and hence the type of

surfactant being used. Particles with large negative or positive zeta

potentials (> 30 mV) will form stable colloidal suspensions since

they tend to repel each other and thereby reduce aggregation.

Table 1

Effect of DMAB concentration on size, PDI and zeta potential (mean s.d., n = 3).

DMAB conc. (% w/v) Size (d. nm) PDI Zeta pot. (mV)

0.01 Aggregation – –

0.025 Aggregation – –

0.05 900 200 0.50 0.08 58 7

0.075 3400 500 0.39 0.09 2 8 *

Fig. 2. DLS size distribution curves and SEM micrographs of blank and drug-loaded

0.10 150 10 0.06 0.02 +57 7

PLGA nanoparticles prepared with the same DMAB concentration (0.10%). Note that

0.125 140 2 0.05 0.02 +75 4

the number of image pixels and signal-to-noise ratio are limited for these SEM

0.15 124 2 0.06 0.01 +80 5 *

images due to a high sensitivity of the nanoparticles to radiation damage. A

minocycline/Ca /AOT-complex at a theoretical D/P ratio of 3% was used.

A.D. Holmkvist et al. / International Journal of Pharmaceutics 499 (2016) 351–357 355

Table 2

for optimizing the drug encapsulation. Table 2 summarizes the

Effect of drug form and D/P ratio on size, PDI, drug content (DC) and entrapment

effect of different combinations on particle size, PDI, drug content

efﬁciency (EE) (mean s.d. n = 3).

and entrapment efﬁciency. All formulations except the 4% HIP-

Drug form D/P ratio (%) Size (d. nm) PDI DC (%) EE (%)

complex, resulted in monodisperse particles (PDI < 0.1) with mean

ﬁ

MC pH 6.4 5 159 4 0.04 0.01 0.06 0.01 1.4 0.2 sizes below 220 nm. The lowest entrapment ef ciencies were

12.5 161 2 0.04 0.02 0.16 0.02 1.5

obtained by the isoelectric approach, most likely because the drug

0.2

is still very water-soluble in the zwitterionic state and diffuses

20 145 4 0.05 0.02 0.26 0.03 1.6 0.2

2 easily into the external aqueous phase during preparation. The

MC/Ca 1 180 5 0.04 0.02 0.11 0.02 12 2

/AOT 1.5 190 10 0.08 0.03 0.27 0.05 20 4 isoelectric formulations were therefore excluded from all further

2 195 1 0.06 0.01 0.82 0.04 43 2 characterization. The highest drug content and entrapment

3 220 6 0.07 0.04 1.12 0.01 43 1

efﬁciency (1.12% and 43%, respectively) were achieved with the

4 700 600 0.4 0.1 – –

HIP-complex in combination with a D/P ratio of 3%. This compares

with others in the ﬁeld who have achieved 1.92% and 29.95%,

respectively, for ion paired PLGA–PEG-dextran sulphate nano-

particles (Kashi et al., 2012) and 9.3% and 46.5%, respectively, for

–

3.2. Drug encapsulation sodium cholate stabilized PEG PLA nanoparticles (Yao et al., 2014).

H NMR spectra of the HIP-complex and control samples are shown

To obtain nanoparticles with high entrapment efﬁciencies by in Fig. 3. Comparison of chromatogram (A) and (B) show changes in

’

the ESD-method, a key parameter is controlling the drugs intensity and shape for minocycline s characteristic peaks at 5.95,

solubility in organic solvents. To increase minocycline’s (MC) 6.83, 7.31, 9.50, 11.53 and 14.87 ppm, which indicates minocycline

lipophilicity, we attempted to replace the hydrochloride salt form complexation with Ca ions. Comparison of chromatogram (C)

’ – with either the isoelectric form (MC pH 6.4) or the HIP-complex and (D) show a broadening of AOT s characteristic peaks at 3.74

2+ 2+

(MC/Ca /AOT). Furthermore, different D/P ratios were evaluated 3.98 ppm in the presence of the minocycline/Ca chelate, strongly

suggesting that a minocycline/Ca /AOT-complex has formed.

Chromatogram (E) of AOT and minocycline in the absences of

Ca appears as a combination of chromatogram (A) and (C),

indicating no interaction occurs between minocycline and AOT in

the absence of Ca ions.

Two different buffering agents were used during particle

preparation, as they are useful in different pH ranges; bis–tris (pH

range 5.8–7.2) was used in the blank and iso-electric formulations

at pH 6.4, and Tris–HCl (pH range 7–9) was used in the HIP-

complex formulations at pH 7.4. The choice of buffering agent

could have effects on particle characteristics such as size and drug

content. The results presented in Tables 1 and 2 however, indicate

that the size increase seen for HIP-complex loaded nanoparticles is

more likely an effect of increased drug content rather than an effect

of using a different buffering agent.

3.3. In vitro drug release

The drug release was quantiﬁed in vitro by a dialysis method.

Unwashed nanoparticles were used both to mimic the future

handling of the formulations, and to prevent morphological

changes and premature drug release, which a washing step using

centrifugal forces might induce (Moreno-Bautista and Tam, 2011).

The release proﬁle of minocycline from minocycline/Ca /AOT-

complex loaded PLGA nanoparticles with a theoretical D/P ratio of

3% (Fig. 4) show an initial burst during the ﬁrst 24 h followed by a

linear release over 30 days. Since unwashed particles were used,

the initial spike in drug release is due to unencapsulated drug. The

subsequent slow release is more likely caused by drug diffusion

and erosion of the polymeric matrix. Cryoprotectants were added

before freeze-drying to prevent the nanoparticles from agglomer-

ating and should be regarded as a part of the formulation. After

resuspension, the initial mannitol and Pluronic F-127 concentra-

tion inside the dialysis bags were approx. 22% and 0.5%,

respectively. Effect of cryoprotectants on minocycline’s diffusion

through the dialysis bags and the release curves are shown in Fig. 4.

The release for only minocycline show that more than 98% of the

free drug was released within two days, which follows the

expected release kinetics from a dialysis bag when replacing the

surrounding release medium at each sampling point and assuming

unhindered diffusion. When cryoprotectants were present how-

1 2+ ever, six days were required for 98% of total drug content release,

Fig. 3. H NMR spectra in CDCl3 at room temperature: (A) MC; (B) MC/Ca ; (C) AOT;

(D) MC/Ca /AOT; (E) MC/AOT. which most likely is an effect of the initial high mannitol

356 A.D. Holmkvist et al. / International Journal of Pharmaceutics 499 (2016) 351–357

et al. (2014) reported a release time of 14 days in PBS (0.01 M, pH

7.4) from their sodium cholate stabilized PEG-PLA nanoparticles

which could be ideal for periodontal disease treatment. Our longer

release proﬁle over 30 days, however, provide a better match with

respect to the normal time course of brain tissue responses

associated with implantation of electrodes (Polikov et al., 2005). It

should be noted that a comparison of in vitro drug-release proﬁles

with others could be misleading, as no standard testing procedures

are used. The in vitro release methods and test conditions are often

chosen with regard to the prospective environment the colloids are

intended for, which may affect the release kinetics. The biological

activity of the released minocycline from our minocycline/Ca

/AOT loaded PLGA nanoparticles remains to be evaluated in vivo.

Prior to clinical use, aspects such as surfactant toxicity and residual

solvent content should also be assessed (Dixit et al., 2015).

4. Conclusion

In this work, minocycline-loaded PLGA nanoparticles were

prepared by a single oil-in-water based preparation technique and

Fig. 4. In vitro release from minocycline/Ca /AOT-complex loaded PLGA nano- characterized for size, drug entrapment and in vitro drug release.

particles with a theoretical D/P ratio of 3% and control curves of minocycline only The inﬂuence of various formulation parameters was investigated

and minocycline together with cryoprotectants.

and optimized with respect to particle size and drug incorporation.

Introduction of a novel HIP-complex of minocycline, Ca and AOT

concentration inside the dialysis bags that causes an increased

resulted in monodisperse particles around 220 nm in size, high

viscosity and thereby a slower intrinsic diffusion. Pluronic F-127 is

encapsulation efﬁciency (43%) and long in vitro drug release

not likely to affect the viscosity at such low concentrations (Gilbert

(>30 days). Our results demonstrate that hydrophobic ion pairing

et al., 1986). The slower intrinsic diffusion is also likely to occur in

with AOT is an effective strategy to incorporate minocycline into

the nanoparticle release study but will presumably level of with

PLGA nanoparticles and could further be applicable to other water-

time as the mannitol molecule is small enough to diffuse through

soluble drugs. The high encapsulation efﬁciency and prolonged

the dialysis tubing. There were no measured differences in

release proﬁle in artiﬁcial CSF suggests that the particles may be

replaced amount of release medium between the three different

useful to accomplish sustained and local delivery of minocycline in

cases and no visual observations of osmotic pressure building up

the CNS. We propose that such nanoparticles delivered from

inside the bags during the release study, which also supports this

implanted electrodes in the brain may be used to reduce the

assumption. The minocycline concentration was evidently higher

foreign body responses and protect the nearby neurones.

inside than outside the dialysis bag during the ﬁrst sampling points

when cryoprotecants were present. This could slow down the Acknowledgements

diffusion of drug molecule from the drug carrier as the

concentration gradient between inside particle and the interior

This work was supported by the Swedish Research Council (No.

of dialysis bag is lower. Due to these diffusional barriers the initial

621-2012-2978), a Linnaeus grant (No. 60012701) from the

drug release may be lower in the generated curve than it would be

Swedish Research Council, The Knut and Alice Wallenberg

if samples ware taken from inside the dialysis bags. This may in

Foundation (No. KAW 2004.0119) and a grant to professor Jens

turn also slow down the overall release from the particles and give

Schouenborg from Vetenskapsrådet (No. K2012-63X-01013-47-4).

rise to a slightly longer release proﬁle. The choice of aCSF as release

The authors would like to thank Dmitry Suyatin for assistance

medium stemmed not only from the prospective application area,

with SEM imaging, Marie Wahlgren, Lars Magnus Bjursten and Lars

the brain, but also for its high content of protective divalent cations

Nilsson for insightful ideas and Lucas Kumosa for writing

2+ 2+

(Mg and Ca ). No degradation of minocycline was detected in assistance.

aCSF during the seven-day stability study (data not shown).

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Paper II

Holmkvist et al. J Nanobiotechnol (2020) 18:27 https://doi.org/10.1186/s12951-020-0585-9 Journal of Nanobiotechnology

RESEARCH Open Access Local delivery of minocycline‑loaded PLGA nanoparticles from gelatin‑coated neural implants attenuates acute brain tissue responses in mice Alexander Dontsios Holmkvist1,2* , Johan Agorelius1,3, Matilde Forni1, Ulf J. Nilsson2, Cecilia Eriksson Linsmeier1 and Jens Schouenborg1,3*

Abstract Background: Neural interfaces often elicit infammatory responses and neuronal loss in the surrounding tissue which adversely afect the function and longevity of the implanted device. Minocycline, an anti-infammatory pharmaceutics with neuroprotective properties, may be used for reducing the acute brain tissue responses after implantation. However, conventional administration routes require high doses which can cause adverse systemic side efects. Therefore, the aim of this study was to develop and evaluate a new drug-delivery-system for local and sustained administration of minocycline in the brain. Methods: Stainless steel needles insulated with Parylene-C were dip-coated with non-crosslinked gelatin and minocycline-loaded PLGA nanoparticles (MC-NPs) were incorporated into the gelatin-coatings by an absorption method and subsequently trapped by drying the gelatin. Parylene-C insulated needles coated only with gelatin were used as controls. The expression of markers for activated microglia (CD68), all microglia (CX3CR1-GFP), reactive astrocytes (GFAP), neurons (NeuN) and all cell nuclei (DAPI) surrounding the implantation sites were quantifed at 3 and 7 days after implantation in mice. Results: MC-NPs were successfully incorporated into gelatin-coatings of neural implants by an absorption method suitable for thermosensitive drug-loads. Immunohistochemical analysis of the in vivo brain tissue responses, showed that MC-NPs signifcantly attenuate the activation of microglial cells without efecting the overall population of microglial cells around the implantation sites. A delayed but signifcant reduction of the astrocytic response was also found in comparison to control implants. No efect on neurons or total cell count was found which may suggest that the MC-NPs are non-toxic to the central nervous system. Conclusions: A novel drug-nanoparticle-delivery-system was developed for neural interfaces and thermosensitive drug-loads. The local delivery of MC-NPs was shown to attenuate the acute brain tissue responses nearby an implant and therefore may be useful for improving biocompatibility of implanted neuro-electronic interfaces. The developed

*Correspondence: [email protected]; Jens.Schouenborg@med. lu.se 1 Neuronano Research Center, Department of Experimental Medical Science, Faculty of Medicine, Lund University, Medicon Village, Building 404 A2, Scheelevägen 2, 223 81 Lund, Sweden Full list of author information is available at the end of the article

© The Author(s) 2020. This article is licensed under a Creative Commons Attribution 4.0 International License, which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons licence, and indicate if changes were made. The images or other third party material in this article are included in the article’s Creative Commons licence, unless indicated otherwise in a credit line to the material. If material is not included in the article’s Creative Commons licence and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder. To view a copy of this licence, visit http://creativecommons.org/licenses/by/4.0/. The Creative Commons Public Domain Dedication waiver (http://creativecommons.org/publicdo- main/zero/1.0/) applies to the data made available in this article, unless otherwise stated in a credit line to the data. Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 2 of 12

drug-delivery-system may potentially also be used for other pharmaceutics to provide highly localized and therefore more specifc efects as compared to systemic administration. Keywords: Neural interface, Minocycline, Gelatin, PLGA, Nanoparticles, Drug-delivery-systems, Biocompatibility, Tissue responses, Immunohistochemistry, Brain

Background biomedical applications, such as drug delivery [19], vac- In order to understand how neuronal networks in the cination [20] and immunomodulation [21]. Tere is mini- brain process information in the awake freely mov- mal systemic toxicity associated with using PLGA even ing animals and how such processing is changed by e.g. in the CNS as the polymer undergoes hydrolysis into lac- learning or diseases, long-term and stable recordings tic acid and glycolic acid, which are eliminated from the of multiple individual neurons are needed. In princi- body through common metabolic pathways [22]. On this ple, this can be achieved by implanting electrodes, often basis, we have previously developed biodegradable mino- referred to as “neural interfaces”, into the neuronal net- cycline-loaded PLGA nanoparticles that provide a sus- works of interest. However, electrodes implanted in the tained release in vitro of minocycline for several weeks brain elicit tissue responses comprising glial activation [23] thus spanning the period of the most intense tissue leading to the formation of a glial scar and loss of nearby infammation after implantation [24, 25]. neurons. Not only do these reactions and loss of neurons Gelatin, a protein-based hydrogel with zwitterionic trigger a reorganization of the networks to be studied, properties, has proven to be a useful biomaterial for con- they also adversely afect the quality of neuronal record- trolled release of several biologically active molecules ings. Considerable eforts have therefore been made to [26]. However, to obtain a sustained drug release from minimize these tissue responses [1]. We have previously a gelatin coating it is necessary to crosslink the gelatin shown that the mechanical properties, such as fexibil- [27] which often involves harsh processing conditions ity, size, anchoring and specifc weight of the electrodes for sensitive molecules like minocycline [28]. Further- are of key importance [2–5]. However, it was only when more, gelatin sufciently crosslinked to ensure sustained combining ultra-fexible mechanical properties with a drug release creates a long-lasting barrier between the gelatin embedding that the ubiquitous loss of nearby implanted electrode and surrounding tissue, which can neurons was, for the frst time, abolished [4]. Moreover, be detrimental for the electrode function [1]. A con- we recently found that gelatin coating promotes healing ceivable approach would thus be to use a non-cross of the blood–brain barrier [6]. Nevertheless, glial cell linked gelatin coating as a carrier for drug-loaded bio- responses remain to some extent in the immediate vicin- degradable nanoparticles. As gelatin is quickly dissolved ity of the electrodes. An alternative and complementary and degraded after implantation, the nanoparticles are approach is to reduce glial responses by coating the elec- released into the adjacent tissue. trodes with anti-infammatory pharmaceutics. Te aims of the present study were to frst develop a A potentially useful pharmaceutic is minocycline, a method for embedding minocycline-loaded PLGA nano- broad-spectrum antibiotic, which also has anti-infam- particles into non-crosslinked gelatin-coatings and then matory, immunomodulatory and neuroprotective efects to clarify whether the addition of the minocycline-loaded [7–11]. It is assumed to act by inhibiting microglial acti- PLGA nanoparticles can mitigate the acute brain tissue vation and their subsequent release of proinfamma- responses. tory cytokines [8, 12, 13] which may be neurotoxic [14]. We report a novel method, suitable for neural inter- Even though minocycline is the most lipophilic anti- faces, to incorporate minocycline-loaded nanoparticles biotic within the tetracyclines, it needs to be adminis- in non-crosslinked gelatin coatings and a signifcant trated repetitively at relatively high doses by conventional reduction in acute microglia activation and astrocytic routes to have efects on the brain due to restricted pas- response nearby the neural implants in mice. sage over the blood–brain barrier [15, 16]. Moreover, minocycline has multiple common adverse side efects Materials and methods [17]. To improve the therapeutic efcacy and to obtain Minocycline‑loaded nanoparticles (MC‑NPs) highly localized efects with less risk of side efects as Te following materials were used for preparing MC- compared to systemic administration, locally delivered NPs: PLGA Resomer RG503H (Sigma-Aldrich, Sweden), drug loaded biodegradable nanoparticles is an attrac- minocycline hydrochloride (Sigma-Aldrich, Sweden), tive option [18]. Poly (d,l-lactic-co-glycolic acid) (PLGA) didodecyldimethylammonium bromide (DMAB) nanoparticles have been widely used for a range of (Sigma-Aldrich, Sweden), sodium bis (2-ethylhexyl) Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 3 of 12

sulfosuccinate (AOT) (Sigma-Aldrich, Sweden). All other triethylamine (2.1 mmol, diluted in DMSO) and AF568 chemicals and solvents used were of analytical grade. Cadaverine (0.44 mL, 0.54 μmol in DMSO) were added. Milli-Q ultra-pure water (Millipore, Sweden) was exclu- Te solution was stirred for 7 h at RT and then the solu- sively used for the preparation of all aqueous solutions. tion was poured into anhydrous diethyl ether to pre- MC-NPs were prepared by a single oil-in-water based cipitate the fuorescent polymer (AF568-PLGA). Te preparation technique previously developed by us [23]. precipitated polymer was dissolved in dichloromethane In brief, PLGA (20 mg) and a hydrophobic ion-pair com- and some drops of a saturated solution of HCl in diethyl plex of minocycline/Ca2+/AOT (3% drug-to-polymer ether was added to neutralize the excess of triethylamine. ratio) was dissolved in an ethyl acetate/methanol mixture AF568-PLGA was fnally precipitated with methanol and (9:1, 1 mL) at room temperature (RT). Te organic phase purifed two more times by dissolution/precipitation was added drop-wise to an aqueous solution contain- using dichloromethane/methanol, respectively. ing a stabilizer (DMAB, 0.10%, 10 −2 M Tris–HCl bufer, F-NPs were prepared according to the particle prepara- pH 7.4, 1.2 mL) under magnetic stirring. Te two-phase tion method described above with the alteration of using system was emulsifed by sonication for 10 min using an AF568-PLGA mixed with PLGA in the ratio of 1:3 (w:w) iced ultra-sonic cleaning bath (VWR USC 300 D, 80W and omitting the minocycline ion-pair complex. Te at 45 kHz). To this emulsion, Tris–HCl bufer (10−2 M, hydrodynamic size, pdi and zeta potential of freshly pre- pH 7.4, 12 mL) was added under constant stirring, which pared F-NPs (before adding cryoprotectants) were deter- resulted in nanoprecipitation. Te suspension was then mined by dynamic light scattering (Zetasizer Nano-ZS, kept in an open beaker with magnetic stirring overnight Malvern Panalytical Ltd., UK), taking the average of three to evaporate the organic solvent. Te suspension was measurements. fnally frozen in liquid nitrogen and lyophilized at − 55 °C and 0.050 mbar for 24 h (Freezone 4.5 model 77,510, Gelatin coating and nanoparticle absorption Labconco, USA). Mannitol (200 mg) and Pluronic F-127 Stainless steel needles (100 μm diameter, Austerlitz (1%, 10−2 M Tris–HCl bufer, pH 7.4, 1 mL) (Sigma- minutiens, Agnthos AB, Sweden) were cleaned in ethanol Aldrich, Sweden) [29, 30] were added as cryoprotectants and insulated with Parylene-C (Galxyl C, Galentis S.r.l., before freeze-drying to prevent the nanoparticles from Italy) (4 μm thick layer) using a compact bench top coat- agglomerating. Tese procedures were identical to those ing system (Labtop 3000, Para Tech Coating Inc., USA). described in previous work in our laboratory [23] and A 30 wt% gelatin (porcine, type A, 300 Bloom, Sigma- result in monodisperse particles 220 ± 6 nm in size, with Aldrich, Sweden) solution was prepared by adding the a polydispersity index (pdi) of 0.07 ± 0.04, zeta potential gelatin to cold HEPES-bufered artifcial cerebrospinal of 55 ± 4 mV, drug content of 1.12 ± 0.01%, entrapment fuid pH 7.3 (aCSF) [31] and allowing the mixture to swell efciency of 43 ± 1% and in vitro drug release character- for 5 min before heating at 60 °C for 1 h under magnetic ized by an initial burst of 20% of the drug load during the stirring. Te gelatin solution was then aliquoted into frst 24 h followed by a sustained release over 30 days. glass vials and stored at 4 °C overnight. Prior to dip-coating, the gelatin solution was reheated and held at 50 °C Fluorescently labelled nanoparticles (F‑NPs) for 30 min to get rid of any air bubbles formed. Te Par- Alexa Fluor 568 (AF568) Cadaverine (Termo-Fisher, ylene-C insulated needles were then dipped 5 mm into Sweden) was conjugated to PLGA through an amide the warm gelatin solution at an immersion and retraction bound between the carboxylic terminus of the PLGA speed of 600 μm/s (dip coating unit model HO-TH-01, and the amine group on the fuorescent dye. First, the Holmarc Opto-Mechatronics Pvt. Ltd., India). Te gela- carboxylic group of PLGA was activated through the tin-coated needles were air-dried for fve min at RT and formation of its N-succinimide ester. Dicyclohexylcar- then stored in a dark enclosure containing silica gel beads bodiimide (11 mg, 53 μmol) and N-hydroxysuccinimide (less than 1% humidity, Type II, Sigma-Aldrich, Sweden). (NHS) (7 mg, 53 μmol) were added to a solution of PLGA Nanoparticles were embedded into the gelatin coating (500 mg, 27 μmol) in anhydrous dioxane (7.5 mL) and by an absorption method. One batch of freeze-dried nan- stirred for 4 h at RT. Precipitated dicyclohexylurea was oparticles (holding 20 mg PLGA and 0.6 mg minocycline) removed by fltration (0.45 µm syringe nylon-membrane was frst resuspended in 1.5 mL water by ultrasonication. flter). Te activated polymer (PLGA-NHS) was pre- Gelatin-coated needles were then immersed 4 mm into cipitated with diethyl ether (10 mL) and then purifed the nanoparticle suspension for 30 s at RT (immersion by dissolving in dioxane and precipitating with diethyl and retraction speed of 600 μm/s), allowing the gelatin to ether three times and fnally dried under reduced pres- swell and absorb the nanoparticle suspension. Te nee- sure. PLGA-NHS (10 mg, 0.54 μmol) was subsequently dles were air-dried for 5 min at RT and then stored in dissolved in dimethyl sulfoxide (DMSO, 1 mL), then a dark dry enclosure (as described above). A schematic Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 4 of 12

illustration of the dip-coating and absorption procedure same setup as for implantation described below) into an is shown in Fig. 1. agarose-flled (0.2 wt % in aCSF) cuvette at RT. Te radial spreading of F-NPs was measured using open source Coating characterization software, FIJI [32]. Te thickness of gelatin-coated needles before and after absorption of MC-NPs was measured at 3 mm distance Animals from the tip using a DS-2MV digital camera (Nikon, 23 transgenic mice (both male and female) that express Japan) mounted on a Nikon eclipse 80i microscope with a green fuorescent protein (GFP) in brain microglia 10× objective. Swelling of the gelatin-coating was quan- (B6129P-Cx3cr1, Jackson laboratories, USA) were used. tifed by measuring the radial expansion over time after All animals recovered from surgery with no apparent applying an aqueous solution containing cryoprotect- adverse efects and all survived for the duration of the ants. Image capture and analysis were performed using experiment. Te mice weighed 34 ± 6 g (mean ± standard the NIS-Elements 3.1 software (Nikon, Japan). deviation) at the time of surgery and they followed a nor- To visualize the absorbed nanoparticles, brightfeld and mal weight curve post-surgery. All mice had free access fuorescent images of F-NP coated needles were acquired to food and water and were kept in a 12-h light/dark using the Nikon eclipse 80i microscope with a ×2 objec- cycle at a constant environmental temperature of 21 °C tive and NIS-Elements software. Te radial distribution and 65% humidity. of F-NPs in the gelatin-coating was imaged using a Zeiss LSM 510 (Carl Zeiss Inc., Germany) scanning confocal Surgery microscope with an Acroplan 40×/0.8W objective. Two Anesthesia was induced by placing the mice in a cham- laser lines, 488 nm and 543 nm were used for sequential ber of 2% isofurane (Baxter Medical AB, Sweden) with excitation of gelatin and F-NPs, respectively. Laser set- 40% oxygen and 60% nitrous oxide. Te head was shaved tings for each laser line was held constant for all images. and subcutaneous injection of local analgesia (Xylocain, Confocal images were collected using the Zeiss LSM 510 3 mg/kg, Lidocain + Adrenaline, Dentsply Ltd., UK and software (version 4.2, Carl Zeiss Inc., Germany). Te Marcain, 1.5 mg/kg, bupivacaine, Aspen Nordic, Den- intensity profle of a vertical cross-section at 50 μm depth mark), was given around skull and neck area. Also 0.5 mL at the 4 mm immersion boarder was quantifed using of saline was given subcutaneously to prevent dehydra- open source software, FIJI [32]. tion. Te mouse was mounted in a stereotactic frame Te in-vitro release of F-NPs from coated needles was (KOPF Instruments, USA). A breathing mask (KOPF imaged using an Olympus IX51 microscope/DP21 digi- Instruments, USA) was attached for constant delivery of tal camera setup (Olympus, Japan) with a 1.4 × objective. anesthesia. Te concentration of isofurane was gradually Te coated needles were inserted (500 μm/s, using the reduced to around 1% during the surgery. To keep the

abc Hydraulic Plunger micropositioner

Glass capillary 5 mm 4 mm 3 mm

50°C22°C Dip-coating Absorption

Gelatin Minocycline-loaded nanoparticles Mineral oil

Fig. 1 Schematic illustrations of dip-coating, nanoparticle absorption and implantation setup. a Dip-coating needle in gelatin (50 °C, down/ up speed 600 μm/s); b absorption of minocycline-loaded nanoparticle suspension (RT, immersion time 30 s, down/up speed 600 μm/s); and c implantation setup with glass capillary and plunger mounted to a hydraulic micropositioner. Needle cut at 3 mm and glass capillary flled with mineral oil to avoid fuid uptake Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 5 of 12

temperature stable a heating pad (3.7 × 14.5 cm, Homeo- 25% sucrose until equilibrated. Subsequently, the needles thermic Monitoring System, Agnthos, Sweden) was used were explanted, the brains were snap-frozen on dry-ice, (body temperature was monitored using a fexible rectal horizontally sectioned (16 µm) on a cryostat (HM560, thermal probe). After disinfection of the skin with 70% Microm, Germany), mounted (serially) onto glass slides ethanol, the skull was exposed with a midline incision. (Super Frost plus, Menzel-Gläser, Germany) and stored Te skin was retracted and the skull was cleaned from at − 20 °C. connective tissue under a stereomicroscope (M651, Leica Te sections were subsequently thawed and rehy- Microsystems, Germany). Craniotomies (∅1 mm) were drated in three 10-min washes of phosphate bufered drilled midways between bregma and lamda, around saline (PBS). Tereafter, incubated in blocking solution, 1 mm laterally of the midline, using a high-speed stere- i.e. 5% normal goat or donkey serum and 0.25% Tri- otaxic drill (Dest 300 IN, model MM 323IN, Silfradent, ton X-100 in PBS (1 h, RT), followed by incubation in Italy). Te underlying dura mater was left intact. Te primary antibodies overnight (blocking solution, RT). holes were rinsed with 0.9% sterile saline solution to clear Primary antibodies included rabbit anti-CD68 (1:1500, away any possible debris and prevent the exposed tissue Abcam Cat#AB125212) to identify activated microglia, from drying. Care was taken to always keep a thin layer rabbit anti-NeuN (1:500, Abcam Cat#AB104225) to of saline on the cortical surface to prevent it from drying. identify neurons, and chicken anti-GFAP (1:1000, Milli- pore Cat#AB5541) to identify astrocytes. After overnight Implantation incubation, slides were washed (3 × 10 min in PBS) and Te needles were prepared for implantation by cut- incubated (2 h, RT) in blocking solution containing sec- ting them to a length of 3 mm and then placed inside a ondary antibodies. Secondary antibodies included don- glass capillary (Microcaps, 1 μL, Drummond Scientifc, key anti-rabbit Alexa Fluor (AF) 647 (1:500, Invitrogen USA). Te glass capillary was mounted onto a hydraulic Cat#A-31573), goat anti-rabbit AF594 (1:500, Invitro- micropositioner (model 2650, KOPF Instruments, USA) gen Cat#A-11037) and goat anti-chicken AF647 (1:500, together with a modifed stainless-steel needle (250 μm Invitrogen Cat#A-21449). In addition, all tissue sec- diameter, Austerlitz, Agnthos AB, Sweden) used as tions were stained with 4′,6-diamidino-2-phenylindole plunger. Te opening of the glass capillary was positioned dilactate (DAPI) (1:1000, Molecular Probes Cat#D3571) over the craniotomy, approximately one millimeter above to visualize cell nuclei. Finally, the slides were washed the brain surface. To avoid water uptake into the capillary (3 × 10 min in PBS) and coverslips were mounted using during implantation and resulting swelling of the gelatin polyvinyl alcohol mounting solution (PVA-DABCO, before entering cortex, the glass capillary was flled with Sigma–Aldrich, Sweden). Te slides were stored in 4 °C parafn oil (Sigma-Aldrich, Sweden). All needles were until imaging. implanted, at a speed of 500 μm/s, to a depth of 3 mm below cortical surface. A schematic illustration of the Image acquisition and analysis implantation procedure is depicted in Fig. 1c. Fluorescent images of stained sections corresponding to a After implantation, a drop of Agarose solution (2% depth of approximately 400–500 µm into the cortex were in saline, type I-B, Sigma Aldrich, Sweden) was placed acquired using a DS-Mv digital camera (Nikon, Japan) on top of the craniotomy and allowed to gel. Te skull mounted on an BX53 microscope (Olympus, Japan) was then sealed using light-cured dental cement (RelyX with a 20 × objective with identical lighting intensity and Unicem, Elipar S10 (430–480 nm), 3M ESPE Inc., Swe- exposure time settings for each respective stain. Regions den). Te skin was closed over the dental cement using of interest (ROIs) were set to 0–50 μm and 50–100 μm surgical clips. For post-surgery analgesia, Temgesic from the border of the implantation site in each of the (0.1 mg/kg, buprenorfn, Indivior UK Ltd, UK) was given photographed sections. subcutaneously. To exclude unspecifc staining in each individual image, intensity thresholds for CD68, CX3CR1-GFP and Immunohistochemistry GFAP were set at 5.2, 7.3 and 3.8 times the background At 3 or 7 days post implantation, the mice were deeply intensity, respectively. Te immunoreactive response anaesthetized by an intraperitoneal injection of pento- for CD68-, CX3CR1-GFP- and GFAP-positive cells was barbital and transcardially perfused with 10 mL of saline, quantifed by measuring the area fraction of the signal followed by 20 mL cold 4% paraformaldehyde (PFA) above the intensity threshold within each ROI. (Termo Scientifc Inc., USA) in 0.1 M phosphate bufer Te number of NeuN-positive cells (with a DAPI- (pH 7.4). Te brains were carefully removed by dissec- positive nucleus as an inclusion criteria) was manually tion and post-fxated in 4% PFA overnight (4 °C). For counted within each ROI and then expressed as num- cryoprotection, the brains were further incubated in ber of neurons divided by the ROI area. Stained cells Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 6 of 12

touching the ROI borders were assigned to the outer ROI to the drug free nanoparticles prepared in our previous on the right side of the image and to the inner ROI on work [23], suggesting that the fuorescently labeled PLGA the left side. DAPI-positive nuclei were manually counted synthesized in this work did not interfere signifcantly in the same manner. Both image acquisition and analysis with the formation of nanoparticles. were done using the NIS-Elements 3.22 software (Nikon, Te gelatin coating stayed intact during F-NP absorp- Japan). tion and drying (Fig. 2a). Te fuorescent intensity profle across the coating was found to be relatively even Statistical analysis (Fig. 2b–d), confrming that the F-NPs were not only Data from 31 successful implantations and subsequent attached to the surface. Furthermore, no visible signs of histological procedures were used for statistics. From agglomeration or clusters of nanoparticles were seen in each implantation site and staining, the mean value of the coating. Te thickness of the coating was similar to available tissue sections in the target depth was used in that of MC-NP coating. Tis, together with similar par- the statistical evaluation. Mann–Whitney test was used ticle properties (size, pdi and zeta potential), indicate to compare the experimental groups within each ROI and that the drug-loaded nanoparticles were absorbed in the time point, p-values < 0.05 were considered signifcant. coatings to the same extent as the fuorescently labeled All analyses were performed using the GraphPad Prism nanoparticles. 8.0.1 software (GraphPad Software Inc., USA). By insertion of the coated needles containing F-NPs in an agarose gel, it was confrmed that the gelatin coat- Results ing stays intact during insertion. An additional movie fle Gelatin coating and nanoparticle absorption shows this in more detail (see Additional fle 1). Moreo- Dip-coating the Parylene-C insulated stainless steel ver, as the gelatin started to swell, the F-NPs nanopar- needles in gelatin resulted in uniform coatings with dry ticles spread radially ∼ 850 μm from the implant over 10 min. thicknesses of 4.8 ± 0.9 μm (mean ± standard deviation, n = 7). Characterization of the prepared minocycline- In vivo efects on the brain tissue response loaded nanoparticles (MC-NPs) including nanoparticle size (220 6 nm), polydispersity index (pdi) (0.07 0.04), Te tissue response to implanted needles coated with gel- ± ± atin containing MC-NPs was evaluated at two diferent zeta potential (55 ± 4 mV), drug content (1.12 ± 0.01%), entrapment efciency (43 ± 1%) and in vitro drug release is described in previous work by us [23] (see also methods). To optimize the absorption of MC-NPs from the a bcGelatin F-NPs suspension, the swelling behavior of the gelatin-coatings was studied. Te gelatin coating swelled 9.8 ± 0.4 (n = 3) times its original thickness after applying the aqueous solution and started to detach from the needle after 43 ± 3 s (n = 3). Terefore, the immersion time was set to 30 s. After absorption of the MC-NP suspension and drying of the gelatin coating, the coating thickness was 9.1 ± 1.2 μm (n = 7), that is almost doubled. Te amounts of nanoparticles and drug absorbed in the coating can be d estimated by considering the concentration of the nano- Gelatin 4 mm F-NPs particle suspension (20 mg PLGA and 0.6 mg minocycline in 1.5 mL water) and volume of the gelatin coating during absorption (85 nL). Te coating was found to swell approximately 10 times its original size during a dipping cycle and, assuming unhindered difusion of the suspension into the gelatin, a total amount of 1 μg nanoparticles and 34 ng minocycline is absorbed in the coating. Fig. 2 Coated implant with absorbed nanoparticles and intensity Fluorescently labelled nanoparticles (F-NPs) were profle. a Brightfeld and fuorescent image of a needle with F-NPs used instead of MC-NPs for further coating characteri- absorbed in the gelatin coating; b, c Two-channel confocal image zation. Dynamic light scattering analysis of freshly pre- (excitation wavelength for gelatin and F-NP was 488 nm and 543 nm, pared F-NPs (before adding cryoprotectants) showed a respectively) centered at the immersion border (box in a) and 50 μm depth; d fuorescent intensity profle of gelatin (green) and F-NPs mean particle diameter of 150 ± 2 nm, pdi of 0.07 ± 0.01 (red) along the arrows in image (b) and (c) and zeta potential of 62 ± 3 mV. Tese results are similar Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 7 of 12

Discussion time points after implantation, 3 days (n = 9) and 7 days We have previously reported that gelatin can be used as (n = 7), and compared to controls with gelatin-coated a coating-material to provide structural support during needles at 3 days (n = 7) and 7 days (n = 8). Te expression of markers for activated microglia (CD68), all micro- implantation of ultrathin fexible electrodes in the brain glia (CX3CR1-GFP), reactive astrocytes (GFAP), neurons [5, 33] and that gelatin itself can signifcantly reduce (NeuN) and all cell nuclei (DAPI) in an inner (0–50 μm microglia activation but not the astrocytic response [4, from border of implant) and an outer (50–100 μm) ROI 6]. Here we developed a novel method for local deliv- surrounding the implantation site was quantifed. Te ery of nanoparticles from gelatin coatings allowing the diameters of the voids did not difer signifcantly between amount of drug to be kept several orders of magnitude the groups (Kruskal–Wallis test). lower as compared to systemic administration. Moreo- ver, the locally delivered minocycline-loaded PLGA Glial response nanoparticles were found to considerably attenuate the acute activation of microglial cells as compared to gelatin At 3 days post implantation, there was a signifcant alone and also cause a delayed reduction of the astrocytic reduction in CD68-positive cells around implants with response. MC-NPs compared to the control in both the inner Te early microglial activation at day 3 and subsequent (p 0.0079) and outer (p 0.0052) ROIs (Figs. 3a and = = astrocytic response on day 7 follows the normal infam- 4a, b). After 7 days (Figs. 3b and 4c, d), the response matory progression after insertion of devices into the had markedly decreased as compared to day 3 for both brain [34, 35]. Microglia begin migrating towards the groups and had almost disappeared in the outer ROI. damaged site after about 12 h [36, 37] whereas astro- However, there was still a signifcant reduction of CD68- cytes start migrating later. Upon microglia activation, positive cells around implants with MC-NPs compared to pro-infammatory cytokines are released which activate the control implants in the inner ROI (p 0.0289). As for = and recruit more microglial cells and appear to induce the CX3CR1-GFP-positive cells (Figs. 3c, d and 4e–h), reactive astrocytes [14]. Conceivably, the substantial there was no signifcant diference between the groups at reduction of activated microglia (CD68-positive cells) either time point. Tese observations suggest that MC- by MC-NP in the present study is coupled to minocy- NPs selectively attenuates the activation of microglial cline’s known inhibitory properties on microglia [8, 13]. cells without efecting the overall population of CX3CR1- Tere was no reduction of the overall microglial cells in GFP-positive cells around the implantation sites. the studied regions of interest, indicating no apparent Te overall astrocytic response after 7 days (Figs. 3f and efect of MC-NP on microglia migration. It remains to be 5c, d) had increased for both groups in both ROIs com- determined, however, if the lowered astrocytic response pared to 3 days (Figs. 3e and 5a, b). However, there was at day 7 found in the present study is a direct efect from a signifcantly lower astrocytic response around implants minocycline or a secondary efect from the decreased with MC-NPs compared to the control in the inner ROI microglia activation. Reduction of both activated micro- (p 0.0401) at 7 days, but not at 3 days. = glia and reactive astrocytes should be benefcial for neuronal survival as these reactive phenotypes are known Neurons and cell nuclei to release components with neurotoxic efects [14, 38]. Te number of NeuN-positive neurons did not signif- Nevertheless, we found no signifcant diference in the cantly difer between groups in either ROI or timepoints number of NeuN-positive cells between the MC-NP and after implantation (Figs. 3g, h and 5e–h). Te density of control groups, suggesting that in this situation there is all cell nuclei (Figs. 3i, j and 4i–l) did not show any sig- no signifcant neuroprotective action of MC-NP. Tis nifcant diferences between groups within ROIs at either fnding may be explained by the used study design with timepoint. Tese and above fndings indicate that the free foating implants and gelatin embedding, both of nanoparticles do not infuence neurons nor total number which known to be benefcial for neuronal survival [39, of cells up to 7 days in mice. 40]. Hence, the need for neuroprotection is less. Te lack

(See fgure on next page.) Fig. 3 Quantitative analysis of the brain tissue responses at 3 and 7 days post implantation. Graphs with activated microglia (CD68) (a, b), CX3CR1-GFP positive microglia (GFP) (c, d) and the astrocytic response (GFAP) (e, f) are presented as the fuorescent area fraction whereas neuronal cells (NeuN) (g, h) and all cell nuclei (DAPI) (i, j) are presented as number of stained cells divided by ROI area. All graphs show the quantifcation of the inner (0–50 μm) and outer (50–100 μm) ROI. Boxes represent the frst and third quartile with median values indicated as horizontal lines within each box and the whiskers show the minimum and maximum values. The horizontal brackets indicate statistically signifcant diferences, p-values are labeled as **(< 0.01) or *(< 0.05) Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 8 of 12

3 days post implantation7 days post implantation

a b

c d

e f

g h

i j Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 9 of 12

3 days post implantation 7 days post implantation Control MC-NPs Control MC-NPs CD68 a b c d

GFP efg h

DAPI ijk l

Merge mn op

Fig. 4 Representative immunofuorescent images of the microglial response. Tissue surrounding gelatin-coated needles with or without embedded MC-NPs at 3 days (two left columns) and 7 days (two right columns) post implantation. Images show activated microglia (CD68) (a–d), CX3CR1-GFP positive microglia (GFP) (e–h), cell nuclei (DAPI) (i–l), and merge (m–p). Scale bar 100 μm = of efect on neurons and overall microglia populations and subsequent enzymatic breakdown into amino acids found here further suggest that the MC-NPs are non- by upregulation of gelatinases (MMP-2 and MMP-9) in toxic. However, there is a need for long term evaluation rodent brain [6]. Te voids in the brain tissue remaining of the biocompatibility of MC-NP. after explanting the implants were of the same dimen- In the present study, we utilized the fact that gelatin sions as the implanted needles, which also suggest that swells but do not dissolve in water at room temperature. the tissue contract tightly around the implant after the Tis made it possible to absorb the particles into the gela- gelatin is dissolved. tin without compromising the thermosensitive minocy- It should be noted that the estimated amount of mino- cline and PLGA nanoparticles. Using non-crosslinked cycline in each implant is less than 10−6 of the amount gelatin as the load-carrying coating material, has previ- usually administered by conventional routes [9, 15], thus ously been shown to preserve the integrity of the neural minimizing risks for unwanted side efects [17]. Further- implant without compromising its function [5]. In the more, a sustained release from nanoparticles eliminates present study, no traces of gelatin were seen in the sur- the need for repetitive drug administration [8]. rounding tissue at either timepoints which is consist- Te in vitro implantation video (see Additional fle 1) ent with known rapid dissolution at body temperature also reveals that the gelatin coating creates a gelatin Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 10 of 12

3 days post implantation 7 days post implantation Control MC-NPs Control MC-NPs GFAP a b c d

NeuN efg h

DAPI ijk l

Merge mn op

Fig. 5 Representative immunofuorescent images of the astrocytic and neuronal response. Tissue surrounding gelatin-coated needles with or without embedded MC-NPs at 3 days (two left columns) and 7 days (two right columns) post implantation. Images show astrocytes (GFAP) (a–d), neurons (NeuN) (e–h), cell nuclei (DAPI) (i–l), and merge (m–p). Scale bar 100 μm =

“track” in which the needle can be freely retracted, thus also be used for other pharmaceutics to provide highly depositing a cylindrically shaped drug reservoir. Tis localized and therefore more specifc efects as com- may be an interesting alternative to injection-based pared to systemic administration. methods as it provides a well-defned local drug release. Supplementary information Supplementary information accompanies this paper at https://doi. Conclusions org/10.1186/s12951-020-0585-9. A novel drug-nanoparticle-delivery-system was devel- Additional fle 1. In vitro implantation in agarose. Schematic illustration oped for neural interfaces and thermosensitive drug- and video of the in vitro implantation of coated needles and spreading of loads. Te local delivery of MC-NPs was shown to fuorescently labeled nanoparticles in agarose gel. attenuate the acute brain tissue responses nearby an implant and therefore may be useful for improving bio- Acknowledgements compatibility of implanted neuro-electronic interfaces. The authors would like to thank Agneta SanMartin and Lina Gällentoft for Te developed drug-delivery-system may potentially assistance with perfusion and immunohistochemistry, and Lucas Kumosa for advice on gelatin coating. Holmkvist et al. J Nanobiotechnol (2020) 18:27 Page 11 of 12

Authors’ contributions neuroprotection and functional recovery after spinal cord injury. Bioma- AH prepared and characterized the nanoparticles and gelatin coatings, terials. 2017;112:62–71. performed perfusion, immunohistochemistry, image acquisition, analysis and 8. Papa S, Caron I, Erba E, Panini N, De Paola M, Mariani A, et al. Early modu- statistical evaluation and gave technical support during surgery. JA performed lation of pro-infammatory microglia by minocycline loaded nanoparti- the surgery and implantations. MF prepared and characterized gelatin coat- cles confers long lasting protection after spinal cord injury. Biomaterials. ings, performed perfusion and image analysis and gave technical support 2016;75:13–24. during surgery. CEL perfused and gave technical support during surgery and 9. Rennaker RL, Miller J, Tang H, Wilson DA. Minocycline increases quality immunohistochemistry. All authors were involved in writing the manuscript. and longevity of chronic neural recordings. J Neural Eng. 2007;4(2):L1–L5. All authors contributed to the experimental design. All authors read and 10. Xue M, Mikliaeva EI, Casha S, Zygun D, Demchuk A, Yong VW. Improv- approved the fnal manuscript. ing outcomes of neuroprotection by minocycline: guides from cell culture and intracerebral hemorrhage in mice. Am J Pathol. Funding 2010;176(3):1193–202. This work was supported by the Swedish Research Council (No. 2016-06195), 11. Zhang Z, Nong J, Zhong Y. Antibacterial, anti-infammatory and neu- NanoLund, Lund University, the Magnus Bergvalls Foundation (No. 2015- roprotective layer-by-layer coatings for neural implants. J Neural Eng. 01176) and the Greta and Johan Kocks Foundation. 2015;12(4):046015. 12. Tikka T, Fiebich BL, Goldsteins G, Keinänen R, Koistinaho J. Minocycline, a Availability of data and materials tetracycline derivative, is neuroprotective against excitotoxicity by inhib- The datasets used and/or analyzed during the current study are available from iting activation and proliferation of microglia. J Neurosci. 2001;21(8):2580. the corresponding author on reasonable request. 13. Kobayashi K, Imagama S, Ohgomori T, Hirano K, Uchimura K, Sakamoto K, et al. Minocycline selectively inhibits M1 polarization of microglia. Cell Ethics approval and consent to participate Death Dis. 2013;4(3):e525. Approval for the animal experiments was obtained in advance from the 14. Liddelow SA, Guttenplan KA, Clarke LE, Bennett FC, Bohlen CJ, Schirmer Malmö/Lund Animal Ethics Committee on Animal Experiments (ethical permit L, et al. Neurotoxic reactive astrocytes are induced by activated microglia. M61-13) and all experiments in this work conform to the regulatory standards Nature. 2017;541:481. of this approval. 15. Fagan SC, Edwards DJ, Borlongan CV, Xu L, Arora A, Feuerstein G, et al. Optimal delivery of minocycline to the brain: implication for human stud- Consent for publication ies of acute neuroprotection. Exp Neurol. 2004;186(2):248–51. Not applicable. 16. Kielian T, Esen N, Liu S, Phulwani NK, Syed MM, Phillips N, et al. Minocy- cline modulates neuroinfammation independently of its antimicrobial Competing interests. activity in Staphylococcus aureus-induced brain abscess. Am J Pathol. Neuronano, A.B., of which JS is a shareholder and board member, holds 2007;171(4):1199–214. patents concerning neural interface technology, including use of gelatin as a 17. Smith K, Leyden JJ. Safety of doxycycline and minocycline: a systematic coating material for drug release and implantation aid. CEL is a co-author of review. Clin Ther. 2005;27(9):1329–42. one patent on neural interface technology including use of gelatin as a coat- 18. Coelho JF, Ferreira PC, Alves P, Cordeiro R, Fonseca AC, Góis JR, et al. ing material for drug release. Drug delivery systems: advanced technologies potentially applicable in personalized treatments. EPMA J. 2010;1(1):164–209. Author details 19. Mir M, Ahmed N, Rehman AU. Recent applications of PLGA based nano- 1 Neuronano Research Center, Department of Experimental Medical Science, structures in drug delivery. Colloids Surf B Biointerfaces. 2017;159:217–31. Faculty of Medicine, Lund University, Medicon Village, Building 404 A2, 20. Silva AL, Soema PC, Slütter B, Ossendorp F, Jiskoot W. PLGA particulate Scheelevägen 2, 223 81 Lund, Sweden. 2 Centre for Analysis and Synthesis, delivery systems for subunit vaccines: linking particle properties to Department of Chemistry, Lund University, Box 124, 221 00 Lund, Sweden. immunogenicity. Hum Vaccin Immunother. 2016;12(4):1056–69. 3 NanoLund, Lund University, Professorsgatan 1, 223 63 Lund, Sweden. 21. 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PLoS ONE. 2015;10(3):0119340. gelatin-based drug carriers by genipin induces changes in drug kinetic 5. Agorelius J, Tsanakalis F, Friberg A, Thorbergsson PT, Pettersson LM, profles in vitro. J Mater Sci Mater Med. 2011;22(1):115–23. Schouenborg J. An array of highly fexible electrodes with a tailored 28. Jain N, Jain GK, Ahmad FJ, Khar RK. Validated stability-indicating densi- confguration locked by gelatin during implantation-initial evaluation in tometric thin-layer chromatography: application to stress degradation cortex cerebri of awake rats. Front Neurosci. 2015;9:331. studies of minocycline. Anal Chim Acta. 2007;599(2):302–9. 6. Kumosa LS, Zetterberg V, Schouenborg J. Gelatin promotes rapid restora- 29. Bardi G, Tognini P, Ciofani G, Rafa V, Costa M, Pizzorusso T. Pluronic-coated tion of the blood brain barrier after acute brain injury. Acta Biomater. carbon nanotubes do not induce degeneration of cortical neurons 2018;65:137–49. in vivo and in vitro. 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31. Protocols CSH. Artifcial cerebrospinal fuid, HEPES-bufered. Cold Spring 38. Block ML, Zecca L, Hong J-S. Microglia-mediated neurotoxicity: uncover- Harbor Protocols. 2011. https://doi.org/10.1101/pdb.rec066696. ing the molecular mechanisms. Nat Rev Neurosci. 2007;8:57. 32. Schindelin J, Arganda-Carreras I, Frise E, Kaynig V, Longair M, Pietzsch 39. Thelin J, Jorntell H, Psouni E, Garwicz M, Schouenborg J, Danielsen N. T, et al. Fiji: an open-source platform for biological-image analysis. Nat Implant size and fxation mode strongly infuence tissue reactions in the Methods. 2012;9(7):676–82. CNS. PLoS ONE. 2011;6:16267. 33. Lind G, Linsmeier CE, Thelin J, Schouenborg J. Gelatine-embedded 40. Köhler P, Wolf A, Ejserholm F, Wallman L, Schouenborg J, Linsmeier CE. electrodes–a novel biocompatible vehicle allowing implantation of Infuence of probe fexibility and gelatin embedding on neuronal density highly fexible microelectrodes. J Neural Eng. 2010;7(4):046005. and glial responses to brain implants. PLoS ONE. 2015;10(3):0119340. 34. Donat CK, Scott G, Gentleman SM, Sastre M. Microglial activation in traumatic brain injury. Front Aging Neurosci. 2017;9:208. 35. Prodanov D, Delbeke J. Mechanical and biological interactions of Publisher’s Note implants with the brain and their impact on implant design. Front Neuro- Springer Nature remains neutral with regard to jurisdictional claims in pub- sci. 2016;10:11. lished maps and institutional afliations. 36. Stence N, Waite M, Dailey ME. Dynamics of microglial activation: a confocal time-lapse analysis in hippocampal slices. Glia. 2001;33(3):256–66. 37. Wellman SM, Kozai TDY. In vivo spatiotemporal dynamics of NG2 glia activity caused by neural electrode implantation. Biomaterials. 2018;164:121–33.

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An implantable microelectrode enabling combined local release of drug loaded nanoparticles and sensing via a liquid contact

Alexander Dontsios Holmkvist1,2, Johan Agorelius1,3, Alle Retelund1,3, Ulf J. Nilsson3, Jens Schouenborg1,3 1 Neuronano Research Center, Department of Experimental Medical Science, Medical Faculty, Lund University, Sweden 2 Centre for Analysis and Synthesis, Department of Chemistry, Lund University, Sweden 3 NanoLund, Lund University, Sweden.

Introduction

Implantable neuro-electronic interfaces that can be used for neuronal recordings and stimulations for long periods of time hold great promise in neuroscience (Wander and Rajesh 2014, Hochberg et al., 2006, Benabid et al., 2009). However, to obtain high quality communications with the highly complex neuronal tissue remains a considerable challenge (Rousche and Normann 1998, Kipke et al., 2008, Schouenborg 2011). Microelectrodes implanted into the brain usually provoke a glial reaction (Edell et al., 1992, Polikov et al., 2005, Szarowski et al., 2003) often thought of as a foreign body reaction, that over time tends to enwrap the microelectrodes and their contacts. As a result, the recordings of neuronal signals get filtered or even blocked (Nicolelis 2003, Kozai et al., 2015) and the electrical impedance and noise are usually increased over the first weeks after implantation (Polikov et al., 2005, Nicolelis 2003,). Moreover, there is invariably a loss of neurons nearby the implanted electrodes (Edell et al 1992, Biran et al., 2005, McConnell et al., 2009,) with the possible exception of ultrathin electrodes (Luan et al., 2017, Wei et al., 2018, Zhao et al., 2019, Musk 2019). To obtain high quality neuro-electronic contact with nearby neurons for long periods of time it is therefore essential to find ways to reduce these tissue responses and avoid loss of neurons.

1 Given that micromotions in the tissue likely contribute to tissue responses (Köhler et al., 2015), there has been a general trend towards ultrathin and highly flexible electrodes that can better follow such movements (Lind et al, 2010, Luan et al., 2017). Another line of strategy aiming at mitigating the glial reactions, has been to coat the microelectrode bodies with anti-inflammatory drugs (Mercanzini et al., 2007, Liu et al., 2013, Lecomte et al., 2015, Holmkvist et al., 2020). However, the long-term recording quality for known microelectrode types still remain relatively poor with respect to spike amplitudes, indicating that the nearest viable neurons are at a considerable distance from the contacts. To obtain high quality neuronal recordings there is thus a need for neuroelectronic interfaces that do not negatively influence the viability of neurons in the immediate vicinity of their contacts. To elucidate how to achieve such interfaces, detailed information on the alterations in the tissue closest to the electrode surface is necessary. Unfortunately, most electrodes used so far must be explanted before histological analysis (Biran et al., 2005), which often results in destruction of the adhering tissue. As a result, information concerning the innermost zone around the electrodes is still very limited. To mitigate some of these problems we have recently developed a new type of microelectrode that is stiff enough to be implanted and then transforms into a flexible and density-matched probe as its interior glucose compartment is dissolved by extracellular fluid entering through an orifice in the thin insulating membrane. Such electrodes have proved to be highly biocompatible, essentially eliminating the loss of nearby neurons, with the possible exception of a few neurons just outside the contact and can be cryosectioned with the electrode still resident in the tissue without being dislocated in the tissue since the fluid inside is also frozen (Agorelius et al., 2020 unpublished). Notably, this construction also allows for local release of biologically active compounds stored inside the insulating membrane through the orifice of the electrode and into the surrounding tissue, thereby offering possibilities to pharmacologically preserve the tissue of interest and also to inhibit migration of glial cells into the electrode. To accomplish sustained (over a month) and localized release of anti-inflammatory drugs, we have previously developed minocycline containing PLGA nanoparticles (Holmkvist et al., 2016) and in a recent study such minocycline containing PLGA nanoparticles embedded in a gelatin coating of microelectrodes were found to substantially mitigate the acute microglia reactions and the astrocytic proliferation around inserted microelectrodes (Holmkvist et al., 2020). In the latter study, we developed a technique to absorb and concentrate the minocycline containing PLGA nanoparticles in the gelatin coating at room temperature by loading in a swollen state and then drying the gelatin with the nanoparticles trapped. The aim of the present work was to enable spatially controlled release of anti- inflammatory nanoparticles to the local milieu just outside the orifice of the membranous microelectrode, with the long term goals of preventing microglial cells to invade and potentially clogg the orifice (i.e. the contact) of the membranous

2 microelectrode and to provide a very useful tool for combined in vivo electrophysiological and histological analysis of local effects of drugs on neuronal tissue with minimal indirect drug effects. To avoid rupture of the thin membrane as water enters its interior through the orifice and the consequent swelling of gelatin, we developed a compartmentalized probe comprising a central unisolated gold lead, surrounded by a first compartment of dry gelatin containing absorbed nanoparticles and a second compartment of dry but readily dissolvable glucose inside the insulating membrane. To characterize the release of nanoparticles through the electrode contact, we used fluorescent nanoparticles (F-NP) and analyzed the kinetics of the release in vitro using time lapse brightfield and fluorescence microscopy. The results show that this construction can provide stable release of nanoparticles, with no signs of nanoparticle clogging within the electrode, over a period of more than an hour.

Methods

Manufacturing of the microelectrodes The first version of the microelectrode tube consisted of an interior gelatin compartment and an insulating Parylene membrane. In short, gold wires (12.5 μm) were first coated with a layer of gelatin (step 1) and then insulated with Parylene-C before an orifice was created in the insulation (step 4). The second version of the tube consisted of a gelatin compartment (step 1) loaded with nanoparticles (step 2) and then additionally coated with an outer compartment of glucose (step 3) before insulated with Parylene C (step 4). Steps 1 - 4 are described in detail below.

Step 1 – Dip coating of gelatin

Preparation of gelatin solution A 30 wt % gelatin solution (porcine, type A, 300 Bloom, Sigma- Aldrich, Sweden) was prepared by mixing gelatin in cold pure milliQ water and allowing the mixture to swell for 5 min before heating at 60 °C for 1 h under magnetic stirring. The solution was then stored at 4 °C overnight and reheated and held at 50 °C for 30 min prior to dip-coating.

Dip coating A frame of gold wires (diameter 12 μm, 99,99%, California Fine Wire company, USA) were prepared by attaching the wires to a custom-made metal frame using double sided tape (7 parallel wires per frame), and the whole frame was then dipped

3 20 mm into warm gelatin solution (50 °C) at an immersion and retraction speed of 600 μm/s and 4 mm/s, respectively (dip coating unit model HO-TH-01, Holmarc Opto-Mechatronics Pvt. Ltd., India). After coating, the wires were air-dried for 5 min at room temperature and then further dried and stored in a dark enclosure containing silica gel beads.

Step 2 – Loading of nanoparticles Preparation and characterization of fluorescent nanoparticles (F-NP) were made in accordance with previous work by our laboratory (Holmkvist et al., 2020), resulting in a mean particle diameter of 150 ± 2 nm, polydispersity index of 0.07 ± 0.01 and zeta potential of 62 ± 3 mV. The nanoparticles were then loaded into the gelatin coated wires by the following absorption method. A batch of freeze-dried F-NPs (holding 20 mg PLGA) was first resuspended in 3 mL water by ultrasonication and 10 μL of the nanoparticle suspension was ejected from a glass capillary attached to a plastic pipette tip using a micro pipette (Microcaps, 1 μL, Drummond Scientific, USA) to form a droplet. Each gelatin coated wire was then pulled through the nanoparticle suspension droplet at a speed of 500 μm/s using a motorized translation stage (MTS50/M-Z8, ThorLabs, USA) (see Figure 1), in effect absorbing the nanoparticle suspension into the gelatin. Finally, the wires were air-dried for 5 min at RT and stored in a dark dry enclosure.

Figure 1. Drug loading achieved by pulling a gelatin coated gold wire through a F-NP suspension droplet hanging from a micropipette.

4 Step 3 - Electrospinning of glucose The wires coated with gelatin and nanoparticles were coated with a layer of glucose by means of electrospinning (see Figure 2). In brief, a solution of glucose in water/ethanol (1 g glucose + 1 g H2O + 0.8 g C2H6O (99.5%)) was ejected form a syringe (0.4 mm) using a syringe pump (ALADDIN programmable syringe pump: 941-371-1003) at a flowrate of 0.4 ml/h. At the same time a voltage of 20 kV (High voltage power supply: HCP 35- 3500 -18498-01-01) was applied between the syringe acting as anode and the wires acting as cathode, separated by a distance of 20 cm. During electrospinning, the frame holding the wires was rotated with a speed of 6 rpm, the temperature was between 45-50 °C, and the relative humidity less than 5 %. After approximately 40 minutes under these conditions the wires were coated with a layer of around 20 μm of glucose.

Figure 2. Schematic illustration of the electrospinning setup, showing the syringe pump ejecting heated glucose solution under influence of high voltage (20 kV) in relation to the rotating fibers being coated, at a distance of 20 cm.

Step 4 - Parylene-C insulation After glucose coating, one layer of Parylene-C (2 μm) was added by vacuum deposition (Compact bench top coating system - Labtop 3000, Para Tech Coating Inc., USA). After this, the wires were cut loose from the frame and coated with a second layer of Parylene-C (2 μm) in order to seal the probe totally. To form the orifice of the tube, a 30x30 μm (mean = 31 x 29, SD = 4.2 x 4.7) (n = 6) hole opening was made through the Parylene coatings at 1.5 mm from the electrode tip using a laser milling machine (Laser mill 50, standard micro-milling system, Nd:YAG laser system, New Wave Research, Inc., Fremont, CA).

5 Characterization of the tube construction

Impedance The orifice in the insulation of the tube where characterized by means of impedance measurements in-vitro. Basically, the opened tubes were immersed into cold saline (0.9 %) and then the impedance was measured at 1 kHz (multielectrode nanoZ impedance tester - 64 channels 1kΩ to ~100MΩ, White Matter LLC, US)

Macrostructure of the tube electrode The macrostructure of the electrode was imaged after each added layer of gelatin using a DS-2MV digital camera (Nicon, Japan) mounted on a Nikon eclipse 80i microscope with a 10× objective. All SEM images were imaged with a Hitachi SU1510 Scanning Electrode Microscope, at 5 KV, using collection of backscattered electrons for contrast.

In vitro release of fluorescent nanoparticles The in-vitro release of F-NPs from the tube electrode was imaged using a DS-MV digital camera (Nikon, Japan) mounted on an BX53 microscope (Olympus, Japan) with a 4x objective. The tube was horizontally inserted into a modified agarose- filled (0.2 wt % in PBS) cuvette at 37 °C. The shift to 0.2% from 0.5%, usually used as brain model, (Pervin et al., 2011) was prompted by the much lower mechanical strength of grey matter of the brain when excluding the mechanically stronger pia mater and arachnoidea (Pomfret et al, 2013). Fluorescent images were acquired (excitation 560/40, emission 630/60, mirror 595 lp) every 2.5 min during a 60- minute recording session. Brightfield images with background illumination of white light were acquired every 5 second for 10 minutes, followed by every 10 second for 20 min, and finally every 30 second for 30 minutes. The wetting of the inside of the tube was quantified by measuring the progression of the waterfront from the hole opening towards the tip at each time frame, using the electrode tip as the reference measuring point. The fluorescent intensity over the same tube area as describe above was also imaged, and it was concluded that the vanishing of the fluorescence level corresponded well to the dissolution of the glucose. Both image acquisition and analysis were done using the NIS-Elements 3.22 software (Nikon, Japan).

6 Results

Tube integrity The first evaluations were made on tubes coated only with gelatin and insulated with Parylene-C (n = 6). However, the integrity of the insulation of these tubes was frequently compromised. After opening the orifice and immersing the tubes in saline, there were stretch marks in the parylene insulation over the entire shaft (n = 6), as seen in SEM images of the parylene layer in Figure 3a. Furthermore, the orifice itself often got enlarged as the edges were cracking (Figure 3b) (n = 6), which suggests that the parylene had been stretched out during the immersion. This is believed to be caused by strong forces caused by the swelling of gelatin when it absorbs water.

A B

Figure 3 a-b. SEM images of pure gelatin tubes after dissolution of the content, showing both (a) cracks or folds in the Parylene sheet, and enlarged hole (b) in the Parylene at the location of the tube opening (n=6).

For this reason, the tubes where henceforth constructed with both an inner compartment of gelatin, loaded with nanoparticles as described above and an additional outer compartment of glucose that dissolves quickly without swelling and leaves room inside the tube for the gelatin to expand into. SEM imaging after the dissolution of the glucose/gelatin tube content indeed showed that there were no cracks in the parylene sheet (Figure 4a) (n = 6) nor any enlargements or rupture of the parylene around the orifice (Figure 4b) (n = 6), after the immersion and dissolution of glucose and gelatin.

7 A B

Figure 4. (a) SEM image of the tube shank after release, showing no cracks or holes in the Parylene-C layer (4 μm) (n = 6). (b) SEM image of the tube opening after the release of content (n = 6).

Macrostructure Dipping the gold wires in gelatin resulted in drop shaped coatings (Figure 5a, d). The average dry drop thickness was larger at the dipping border compared to 3.5 mm distance from the border, 22.2 ± 2.2 μm (mean ± standard deviation, n = 9) and 19.3 ± 1.7 μm respectively (table 1). Absorption of the F-NP suspension (Figure 5b, e) increased the dry drop thickness slightly, but retained a wavy structure. Electrospinning of glucose and coating with one layer of Parylene-C therefore resulted in a wavy structure (Figure 5c, f).

Table 1. Average coating thickness after each added layer at the gelatin dip border and at 3.5 mm distance from the border (n = 9).

Layer Thickness at border (μm) Thickness at 3.5 mm (μm) Gelatin 22.2 ± 2.2 19.3 ± 1.7 Gelatin + F-NPs 22.7 ± 2.3 19.7 ± 2.2 Gelatin + F-NPs + Glucose 44.4 ± 2.6 45.7 ± 4.4

8 Gelatin + Fluorescent-NPs + Glucose a b c Border

d e f 3.5 mm

100 μm

Figure 5. Brightfield and fluorescent images of a gold wire after coating with gelatin (a, d); absorption of F-NPs (b, e); and electrospinning of glucose with one layer of Parylene C (c, f). The images were taken at the gelatin dip border (a- c) and at 3.5 mm distance from the dip border (d-f).

Tube impedance The tube, with its orifice that allows extracellular fluid to enter the tube so that the glucose is dissolved, showed a quick decrease in impedance after immersion into saline, stabilizing at 110 ± 20 kOhm (n = 5) at 1 kHz, which is suitable for electrophysiological recordings.

In vitro release of fluorescent nanoparticles The release of the tube content (n = 3) was indicated as the level of glucose dissolved measured as the progression of the water front inside the tube, a level that overlapped well with the loss of fluorescence in the tube (Figure 8), indicating the release of F-NPs. The release rate decreased rapidly during the first 15 min and then more slowly over the following 30 minutes, as illustrated by the total volume dissolved over time in Figure 6 and the release rate over time in Figure 7. No clogging at the orifice was seen.

Mean Release from tube (ml) 10

Volume (ml) Volume 2

0 -5 5 15 25 35 45 time (min)

Figure 6. The dissolved volume from inside the tube as a function of time, mean and standard deviation.

Mean release rate (ml/min) 0,7 0,6 0,5 0,4 0,3 0,2 0,1 0

-5e(ml/min) rat Release 5 15 25 35 45 time (min)

Figure 7. Release rate of dissolving tube (mm3/min) as a function of time, mean and standard deviation.

10 1 min 2 min 5 min 10 min 30 min

Tube opening

Figure 8. Image series of the in vitro release of fluorescently labeled nanoparticles from the tubular microelectrode in agarose.

Over the course of approximately 1 hour, tubes were emptied up to 1.5 mm from the orifice in both directions, corresponding to a dissolution of a total amount of 0.2 nL of gelatin loaded with nanoparticles. Assuming similar nanoparticle- concentration inside the gelatin as in the bulk solution used for absorption and a swelling ratio of the gelatin of around 4.6 x, this would correspond to the release of 19 ng of nanoparticles.

11 Discussion

The development of the device described here is motivated by the strong need for improved means of communication with viable neurons and neuronal networks under physiological conditions (Polikov et al., 2005, Kipke et al., 2008, Schouenborg 2011). Existing implantable electrodes are usually hampered by glia cells adhering to the probes (Edell et al., 1992, Biran et al., 2005) including the very contacts, following implantation, thus causing deterioration of the recorded signals (Rousche and Normann, 1998, Salantino et al., 2017). In parallel, ongoing studies we have found the novel tube electrode to be outstanding with respect to biocompatibility/recording quality (Agorelius et al., 2020 unpublished) and microglia do not seal the tube orifice. However, it appears that microglia can migrate into the tube as ED1 and DAPI stained cells can be found eventually. Also, remnants of astrocytic GFAP stained cells have been observed inside the tubes. To reduce the risk of clogging, release of pharmaceutics from the interior of the tube that reduce migration of glial cells into the tube is an interesting possibility. In the present work, we took the first steps towards enabling this possibility and developed an implantable and highly biocompatible device that allows release of anti- inflammatory drugs and dynamic recordings of neuronal activity adjacent to the tube orifice. Notably, the device also allows refined pharmacological studies with less confounding factors and side effects than for example after systemic administration of drugs (Fagan et al., 2004, Smith et al., 2005).

Technical considerations We made use of and modified a novel implantable electrode construction that, once implanted, transforms into a flexible tube. The tube comprises a lead with a large non-isolated surface, dissolvable materials (gelatin and glucose) that are stiff during insertion and thus provide structural support and a 4 µm thick Parylene-C electrically insulating membrane. The dissolvable material harbors nanoparticles that can be loaded with drugs. The coat has one opening in the distal part of the probe, allowing ionic and fluidic contact with the surrounding tissue. Figure 9 provides a schematic of the developed tube electrodes loaded with drugs and dynamics of the release of drug containing nanoparticles. The time lapse fluorescence microscopy (Figure. 8) confirmed the general idea, but the F-NPs are too small to be seen individually with fluorescence microscope once released.

Figure 9. Schematic illustration of tube electrode composition (a) and proposed nanoparticle release (b-c) after implantation; (b) Water from the surrounding enters the tube and causes gelatin swelling and release of nanoparticles inside the tube; (c) As the tube continues to be filled with water, the gelatin coating dissolves and nanoparticles are released into the surrounding tissue.

Loading capacity Although not directly evaluated, it appears that the drug load can be modified in many ways. For example, the amounts of nanoparticles loaded in the gelatin is a function of the gelatin swelling during NP loading (Holmkvist et al., 2020). In the present study, utilizing NP loading during a phase when the gelatin coated wire is dragged through a droplet with NP in suspension, the speed of the wire through the droplet will determine the extent to which the gelatin is wetted. This in turn determines the swelling and thus the amounts of nanoparticles trapped in the gelatin when drying. The total amount of particles loaded this way inside the tube will of course also depend on the total amount of gelatin and concentration of nanoparticles in the suspension. The drug load used here was calculated to be similar to that known to show effects on microglia and astrocytes (Holmkvist et al., 2020).

Nanoparticle release characteristics The present study showed that the force produced by gelatin inside the tube when wetted can be strong enough to rupture the Parylene coat and even to stretch out and deform the orifice of the tube. By introducing quickly dissolving glucose, as an outer layer on top of the gelatin, thereby creating a space in which gelatin can swell, this problem was effectively mitigated. This arrangement also has implications for the kinetics of the release. In our previous study, it was clear that as gelatin takes up

13 water and swells, NP are freed from the gelatin matrix and quickly diffuse away. When inside a tube, the swelling of gelatin may also significantly push out NP through the hole in the initial phase. From the release curves there are indications of two phases – an initial phase lasting around 15 min during which the release of NP drops rather fast followed by a phase with slowly declining NP release. Presumably, the shift from fast to slow decline in release corresponds to the time course of gelatin swelling and to some extent progressive increase in diffusion time to the hole of the tube. Gelatin swelling/dissolution are parameters that can be manipulated by for example changing the degree and type of cross-linking and can thus be anticipated to affect NP release. Taken together, there are therefore ample possibilities to tune the release rate as well as the drug load. The size and surface charge of the nanoparticles can greatly influence their behavior in vivo (Popović et al., 2010, Hoshyar et al., 2016) due to factors such as diffusion and cellular uptake. For example, particles smaller than about 40 nm redistribute in the tissue mainly through mechanisms of diffusion (Hayes and Kruger 2014), whereas larger particles are less motile (Eugin et al., 2017). This suggest that the relatively large particles used here (~200 nm) will be relatively stationary in the tissue, thus the released drug will be concentrated to the region outside the tube orifice, providing for potentially very long-lasting effects (month) in the region where it is most needed (Holmkvist et al., 2016). With this electrode and NP-enabled sustained release available, opportunities are open to evaluate how different drugs, by themselves or in various combinations affect astrocyte activation, microglia differentiation and activation, and neuron viability/activity. For example, effects of Minocycline on microglia activation (Holmkvist et al., 2020) in the acute inflammatory stage could be compared with or combined with the effects of drugs with potential to regulate profibrotic type 2 microglia active in chronic inflammation and fibrosis (Boza-Serrano et al., 2019) near the electrode. Furthermore, effects of other drugs such as neurotrophic factors, antibiotics etc. (co-encapsulated or in separate NP) in combination with microglia or astrocyte-regulating drugs can be evaluated and fine-tuned with the microelectrode design described herein.

Conclusion

We have developed a tubular recording/stimulating neuroelectronic interface that can harbor and release nanoparticles to the tissue of interest, with the primary goal to provide a clog-free tubular electrode and also to control inflammatory reactions in the tissue. Drug load and release rates can be tailored and need to be optimized in future in vivo studies.

14 References

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ALEXANDER DONTSIOS HOLMKVIST Nanoparticle-based drug delivery systems for neural interfaces 2020:120

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ISSN 1652-8220 ISBN 978-91-7619-983-1 Lund University, Faculty of Medicine Doctoral Dissertation Series 2020:120 Department of Experimental Medical Science Medical Experimental of Department