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Nano Energy

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Communication

Flexible wireless powered drug delivery system for targeted administrationon cerebral cortex

Sang Hyun Sunga,1, Young Soo Kimb,1, Daniel J. Joea,1, Beom Ho Muna, Byoung Kuk Youa,Do Hee Keumc, Sei Kwang Hahnc, Magnus Berggrend, Daesoo Kimb,⁎, Keon Jae Leea,⁎

a Department of Materials Science and Engineering, Korea Advanced Institute of Science and Technology (KAIST), 291 Daehak-ro, Yuseong-gu, Daejeon 34141, Republicof KoreabDepartment of Biological Sciences, Korea Advanced Institute of Science and Technology (KAIST), 291 Daehak-ro, Yuseong-gu, Daejeon 34141, Republic of Koreac Department of Materials Science and Engineering, Pohang University of Science and Technology (POSTECH), 77 Cheongam-ro, Nam-gu, Pohang 790-784, Koread Laboratory of Organic Electronics, Department of Science and Technology, Linköping University, SE-601 74 Norrköping, Sweden

A R T I C L E I N F O

Keywords:Drug deliveryFlexible deviceImplantable deviceWireless power transfer

A B S T R A C T

The controlled drug delivery devices helps timely drug administrations and maintenance of effective dose tomaximize curing effects with minimal side effects. Application of this technology to various body parts has beenlimited, especially in organs with curved surface, such as the brain and the eye. Herein, we report a flexible drugdelivery microdevice (f-DDM) for controlled administration on the curved organ surface. The unique structure ofthe f-DDM consists of freestanding gold membranes over the multireservoir array was implemented by reversingthe typical fabrication order of the reservoir and sealing membrane. We optimized the design of the f-DDM by afinite element analysis to prevent thermal damage during the laser transfer and the applying current density forreliable drug release through an electrochemical analysis. The wireless power transfer system was applied to f-DDM, which shows stable wirelessly powered operation. The f-DDM was flexible enough to be implantable onthe curved cerebral cortex and successfully adopted for delivery of two different chemicals or prevention ofseizure activity using an anti-epileptic drug. Our study opens a new avenue for the controlled, region-specific,and combinatorial application of drugs, the key factors for precision medicine.

1. Introduction

The development of flexible electronics is opening a new avenuetoward the era of wearable and implantable biomedical devices in-cluding blood pressure sensors, microneedle patches, smart contactlenses, and optogenetic devices [1–6]. Flexible biomedical micro-devices, which are small, cost-effective, and capable of health mon-itoring in daily life, can be conformally implanted in various locationsincluding skin, eyes, and other organs providing physiological in-formation in real time [7,8]. In particular, application of these flexiblebiomedical devices has recently been expanded to the area of drugdelivery. A drug delivery system that operates according to blood glu-cose level was demonstrated to treat diabetes by drug administrationfrom thermo-responsive microneedles in response to feedback signals ofa glucose sensor [9].

The cerebral cortex is the largest brain area and consists of func-tional domains along a curved structure. Dysfunction of a specificcortical domain has been associated with neurological disorders such as

ischemia, epilepsy, and dementia [10–12]. The application of flexibledevices has been proposed for searching disease-related areas and tar-geted drug delivery in the cortex [13–15]. Previously, our group re-ported a flexible electrocorticography (ECoG) system for optogeneticmapping, which was implanted through a small cranial slit withoutbrain damage [16].

Although the flexible ECoG optogenetic system establishes a con-ceptual basis for sensing disease-related cortical domain, targeted drugdelivery has remained as a challenging issue. The targeted drug de-livery is the key to treat cortical dysfunction because drug diffusion toother cortical areas results in unexpected side effects [17,18]. Whilecontrolled-release microchips with the precise control of drug releaseand long-term operation for subcutaneous administration has beendemonstrated in human, these therapeutic microdevices are difficult tobe conformally implanted onto the brain surface due to their bulky sizeand lack of flexibility [19,20].

https://doi.org/10.1016/j.nanoen.2018.06.015Received 12 May 2018; Received in revised form 5 June 2018; Accepted 5 June 2018

⁎ Corresponding authors.

1 These authors contributed equally to this work.E-mail addresses: daesoo@kaist.ac.kr (D. Kim), keonlee@kaist.ac.kr (K.J. Lee).

Nano Energy 51 (2018) 102–112

Available online 07 June 20182211-2855/ © 2018 Elsevier Ltd. All rights reserved.

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2. Experimental section

2.1. Animals

Animal care and experiments were conducted in accordance withthe guidelines of the Institutional Animal Care and Use Committee ofKorea Advanced Institute of Science and Technology (KAIST). All micewere maintained on a 12-h light/dark cycle (light cycle beginning at6:00 A.M.) at a temperature of 23 °C. Food and water were supplied adlibitum. All animal experiments employed 3–6 month-old male mice.

2.2. Device fabrication

A 50 nm hydrogenated amorphous silicon (a-Si:H) exfoliation layerand a 1.5 μm-thick SiO2 buffer layer were deposited by plasma en-hanced chemical vapor deposition (PECVD) on a rigid glass substrate.Ti (20 nm)/Au (100 nm)/Ti (20 nm) layers were deposited and pat-terned by an electron beam evaporator and conventional lithographyprocess for microfabrication. The multireservoir array was patterned bySU-8 photoresist. The device was bonded to a 25 μm-thick flexible PETsubstrate, followed by the ILLO process to remove the rigid substrate.After the exfoliation, the device was passivated by SU-8, except for thedrug releasing membrane. Lastly, the SiO2 buffer layer and the Ti layerwere wet-etched to open the electrodes.

2.3. Heat transfer simulation

The simulation was based on COMSOL Multiphysics, using time-dependent study of application mode Heat Transfer in Solids (ht). Thematerial parameters are based on the COMSOL database: a-Si (Thermalconductivity, k= 1.3W/(m K). Density, ρ =2259 kg/m3. Heat capa-city, Cp = 700 J/(kg K)), SiO2 (k= 1.4W/(m K), ρ =2200 kg/m3, Cp

= 900 J/(kg K)), Si3N4 (k= 20W/(m K), ρ =3100 kg/m3, Cp = 700 J/(kg K)), and Silica glass (k= 1.38W/(m K), ρ =2203 kg/m3, Cp

= 703 J/(kg K)). Laser heating by a laser beam with an energy densityof 512mJ/cm2 and a time pulse of 60 ns was implemented by applyingthe Beer-Lambert law.

2.4. In vitro electrochemical analysis

Electrochemical characteristics of the f-DDM were recorded in 1Xphosphate-buffered saline (PBS) at room temperature using a Keithley4200-SCS under constant DC voltage or current bias. An Ag/AgCl re-ference was used for three terminal measurement. Linear potentialsweep was applied from 0 V to 2 V with respect to the Ag/AgCl re-ference electrode in a PBS medium. Time-lapse of images of f-DDMwere taken by optical microscope.

2.5. Brain implantation in live mice

All surgical apparatuses were sterilized with 70% ethanol solution,excepting the f-DDM device sterilized with 70% isopropyl alcohol. Forimplantation surgery of the f-DDM, mice were initially anesthetizedwith avertin (20mg/mL of tribromoethanol, 20 μL/g i.p.) and fixed on astereotaxic apparatus (David Kopf Instruments, Germany). Aftermaking a cranial window (1.0×4.0mm) on the target brain region,the f-DDM was slowly slid beneath the skull into a meningeal space,then bonded on the skull using dental cement (Super Bond, SunMedical, Japan). After recovery, electrical stimulation was applied witha stimulation isolator (World Precision Instruments, USA). The micewere sacrificed three days after stimulation, and the brain was preparedwith 4% formaldehyde. The brain was sliced with the thickness of50 µm by a Vibratome (Leica, Germany).

2.6. Fluorescence imaging

Three different fluorescent materials were used for thebrain administration: 10 mg/mL of 1,1’-dioctadecyl-3,3,3’,3’-tetra-methylindodicarbocyanine, 4-chlorobenzenesulfonate salt (DiD, λmax,

Em = 665 nm) (D7757, Molecular Probes, USA), 10 mg/mL of 1,1’-dioctadecyl-3,3,3’,3’-tetramethylindocarbocyanine perchlorate (DiI,λmax, Em = 549 nm) (No. 468495, Sigma-aldrich, USA), and 2 mg/mLof BODIPY TMR-X-conjugated Muscimol (M23400, Life technologies,USA). To verify minimal damage to the cerebral cortex after the deviceimplantation and drug administration, extra fluorescence images wastaken with 10mL of 4,6-diamidino-2-phenylindole (DAPI, Vectashieldantifade mounting medium with DAPI) (H-1200; Vector laboratories,USA). Microscopic images were captured with a fluorescent micro-scopy (IX51, Olympus, Japan; LSM780, Carl Zeiss, Germany), and 3Dimage reconstruction was performed with Free-D software (http://free-d.versailles.inra.fr) from 2D fluorescent images of sequentialbrain slices.

2.7. Statistics

All data are presented as mean ± SD. Statistical significance ofbending test was analyzed by one-way analysis of variance (ANOVA).Sample size was n=4, and the p value was 0.799 and 0.710, indicatingthat the bending test does not have significant effect on device opera-tion. Statistical analysis was carried out with Microsoft Excel.

3. Results

3.1. Fabrication of flexible drug delivery microdevice

Fig. 1a illustrates the structure of the f-DDM composed of a poly-ethylene terephthalate (PET) substrate, a multireservoir array, a metalsealing membrane, and a passivation layer. The f-DDM was implantedonto the cerebral cortex through a small cranial slit as shown in Fig. 1b.The 25 μm-thick PET substrate was specifically selected for the f-DDMto provide appropriate flexibility and handling for the sliding insertionto the subdural space as demonstrated in our previous study [16]. The300× 300 µm2-sized microreservoir layer (with the thickness of 15 µmand the volume per reservoir of 1.35 nL) was patterned using epoxy fordrug storage. The specification of f-DDM in this study was determinedfor localized treatments in a mouse brain, which can be flexibly ad-justed for a subject, drug, and target region. The metal layer was used asboth electrodes and a sealing membrane on top of the microreservoir.We chose gold as the metal sealing layer due to its excellent chemicalresistance and biocompatibility, which can offer long-term in vivo op-eration of the f-DDM. Finally, the f-DDM was passivated by an addi-tional epoxy layer to prevent undesired reaction and leakage currentoutside of the sealing membrane. All of these materials used for thefabrication of the f-DDM are known to be biocompatible for implanteddevices as reported in numerous studies [21–24].

The detailed fabrication process of the f-DDM is shown in Fig. 1cand S1. First, an exfoliation layer of hydrogenated amorphous silicon(a-Si:H) followed by a 1.5 μm-thick SiO2 buffer layer were deposited ona glass substrate using plasma-enhanced chemical vapor deposition(PECVD). The SiO2 layer was used to protect the reservoir from thermaldamage during the ILLO process [25]. In order to fabricate a free-standing gold membrane that sits on top of the microreservoirs, a metalelectrode was deposited using an e-beam evaporator prior to the mi-croreservoir formation. The reservoir layer was formed by an epoxy-based photoresist on top of the patterned metal electrode. The micro-reservoir was then filled with the drug in solid (~ 1.35 nL) and sub-sequently bonded to a flexible PET substrate with UV adhesive. Aftercuring the adhesive, the f-DDM was flipped over and detached from theglass substrate via laser irradiation. Another epoxy layer was coated topassivate the f-DDM, and it was subsequently patterned to expose the

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Fig. 1. Structure and fabrication process of flexible drug delivery microdevice (f-DDM). (a) Layer components of f-DDM composed of a PET substrate, an epoxymicroreservoir, a metal membrane, and a passivation layer. (b) Schematic of the f-DDM inserted beneath the skull in a live mouse through a small cranial slit. The f-DDM was attached on the mouse skull by dental cement after the implantation. (c) Fabrication process of the f-DDM. (i) Reservoir/electrode patterning and druginjection on top of the rigid glass substrate. (ii) Transfer to the flexible PET substrate by inorganic laser lift-off (ILLO). (iii) Fabricated f-DDM after the passivationlayer patterning process. (d) Cross sectional schematic and (e) SEM image of f-DDM showing the freestanding gold membrane over the flexible microreservoir. (f)Photograph of f-DDM mounted on a 7mm diameter glass rod. (g) Magnified optical images before and after applying the electric bias that indicate dissolution of thegold membrane.

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sealing membrane for an electrically-controlled electrochemical reac-tion. The unique structure of the f-DDM consisting of freestanding goldmembranes over the multireservoirs was implemented by reversingcommon fabrication sequences of the reservoir and sealing membrane.

The direct formation of a thin metal layer over the reservoirs ishighly challenging because of the step coverage of deposited metalfilms and undesirable contamination of the drug. The cross-sectionalschematic illustration of the microreservoir in Fig. 1d shows that thedrug-loaded reservoir is encapsulated by the gold sealing membrane toprevent any undesired release of the drug. The presence of the free-standing gold membrane can be observed in the scanning electronmicroscope (SEM) image in Fig. 1e. The SEM image clearly exhibitseach layer component including the freestanding gold sealing mem-brane directly above the drug-loaded microreservoir. Photographs ofthe fully fabricated device conformally mounted on a 7mm diameterglass rod are shown in Fig. 1f, demonstrating the high flexibility of thef-DDM. The gold sealing membrane becomes electrochemically dis-solved by the external electric bias, which leads to drug release from thereservoirs. The magnified optical images of the device in Fig. 1g in-dicate that the gold membranes completely disappeared from the fivereservoirs after applying constant DC currents. This total dissolution ofthe gold membranes confirms that the metal electrodes were not da-maged during the ILLO transfer.

3.2. Simulation of ILLO transfer process

In order to optimize the structure and design of the f-DDM, we in-vestigated the FEA of laser-induced thermal annealing during the ILLOprocess. The ILLO process was based on the melting of an opticallyreactive exfoliation layer by a focused laser beam [26]. During the ILLOprocess, a XeCl laser beam with energy density of 512mJ/cm2 (2mJper 625× 625 µm2) is irradiated to the a-Si:H exfoliation layer within atime pulse of 60 ns. A buffer layer is required to protect the drug in themicroreservoirs and metal electrodes against an excessive amount ofthermal energy, which is diffused from the exfoliation layer. COMSOLMultiphysics was used to specifically determine the SiO2 buffer layerthickness by analyzing the heat transfer in the f-DDM during the laserirradiation. The laser-induced heating model was implemented by theBeer-Lambert law as follows [27–29]:

= − = = − −Q dI zdz

αI z α R I αz(z)( )

( ) (1 ) exp ( )0

where Q is the laser-induced heat source, I is the laser intensity, α isthe absorption coefficient, and R is the reflectance. The temperature ofthe a-Si:H layer instantly rises above 2500 K, as shown in Fig. 2a. The a-Si:H layer melts at around 1400 K, leading to delamination of thetemporary glass substrate [30,31]. The time dependent temperatureprofile in Fig. 2b shows that the temperature of the exfoliation layer

Fig. 2. Time-lapse images of f-DDM after applying the electrical bias. (a) FEA results of heat transfer during the ILLO process. The laser-induced heat was im-plemented according to the Beer-Lambert law. (b) Temperature profile as a function of position in the z-direction. The maximum temperature of the a-Si exfoliationlayer was 2711 K with an irradiated laser energy density of 512mJ/cm2. The temperature instantly rises up to the maximum value at 120 ns, and diminishes after9 μs. (c–i) Magnified optical and fluorescence images of f-DDM at different time intervals from 10 s to 130 s (j) The voltage-time curve with applied current density of1.1 mA/cm2. The alphabet letters the points consistent with the optical images of Fig. 2c–i. The freestanding gold membrane is fully dissolved at 130 s as indicated bythe rapid transition in the voltage-time curve, which is also demonstrated by the fluorescent image in the insets of Fig. 2c and i.

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rises up to maximum temperature of 2711 K at 120 ns (the laser heatingstarts at 60 ns, and ends at 120 ns). After 120 ns, the generated heatdecays in the confined area, and eventually diminishes at 9 μs. The timedependent simulation results verify that the 1.5 μm-thick SiO2 bufferlayer can sufficiently protect the device layer from the laser-inducedheat. Furthermore, the ILLO process can prevent mechanical damagesof the device such as deformation, crack, and expansion of the de-posited film. While a polymer sacrificial substrate is vulnerable totemperature and external stress, the ILLO process is based on a rigidinorganic substrate that can provide stable environments for the fab-rication process of a freestanding gold membrane such as e-beam eva-poration and photolithography. Since the ILLO process does not induceany degradation of the device, the biocompatibility of the materials canbe maintained after the device fabrication processes. The biocompat-ibility of the biomedical device fabricated via ILLO process had beenreported in previous studies [32,33].

To verify the simulation-based design, the in vitro performance of f-DDM was characterized as shown in Fig. 2c–j and Video S1. The goldsealing membrane was designed to be dissolved as AuCl4− ions in achloride solution by anodic corrosion with external electric bias[34,35]. The device was operated by applied electrical current densityof 1.1 mA/cm2 (i.e. 1 μA at 300×300 µm2-sized gold membrane) in1× phosphate-buffered saline (PBS) solution, whose osmolality and pHvalue are similar to those of human body fluid. The magnified opticalimages of Fig. 2c–i explicitly show the time-lapse disappearance of goldmembrane covering the microreservoir in real time. The gold film wascompletely dissolved in the PBS after 130 s, which was confirmed by thedetection of vivid green fluorescence inside the microreservoir (seeinset of Fig. 2i). This in vitro experiment clearly demonstrates that the f-DDM was successfully fabricated and operates properly in the PBS en-vironment. Each letter in the voltage vs. time curve of Fig. 2j corre-sponds with each optical image of Fig. 2c–i. Constant voltage wasmeasured in the dissolving region (Fig. 2c–h), while a rapid transitionto higher voltage was measured in the fully dissolved region (Fig. 2i),indicating complete dissolution of the gold membrane.

Supplementary material related to this article can be found online athttp://dx.doi.org/10.1016/j.nanoen.2018.06.015.

3.3. Wirelessly powered operation of f-DDM

The integration of appropriate power source instead of external power isimportant for the clinical application of f-DDM. Wireless power transfer ispromising candidate since the f-DDM requires input power of 1.34 µW forthe dissolution of one gold membrane, which is small enough to be wire-lessly powered [36–38]. The flexible wireless power transfer system wasestablished as shown in Fig. 3a. The AC voltage source (Agilent 33,250A)generating radio frequency (RF) power with 5V amplitude and 13.56MHzfrequency (ISM band) was connected to flexible antenna (transmitter coil)for the wireless power transmission in near-field range. The generatedelectrical energy was transferred to another flexible antenna (receiver coil)by inductive coupling. The rectifying circuit composed of voltage regulatorand peak rectifier converted the transmitted AC power to DC power. Thestable DC voltage of 3.29V was measured after the wireless transfer andrectification with input AC voltage of 5V.

Fig. 3b demonstrates the change in efficiency of flexible wireless powertransfer system with increasing distance or bending radius. To investigatethe effect of distance between two flexible antennas, the input power wasfixed to 0.5W (5V with 50Ω) and the distance was increased from 0.5 cmto 2.0 cmwhile both antennas were in flat state. As shown in the Fig. 3b, thepower transfer efficiency linearly decreased with the increase of distancefrom 1.22% (output voltage of 4.28V with 3000Ω) to 0.06% (output vol-tage of 0.96V with 3000Ω). On the other hand, the effect of bending radiusof receiver antenna to power transfer efficiency was examined by fixing thetransmitter coil to receiver coil distance as 0.5 cm. The power transfer ef-ficiency was stable with the bending radius of 3.0 cm and 2.0 cm, and de-creased with larger bending radius.

The f-DDM was connected to the wireless power transfer system todemonstrate the wirelessly powered operation of f-DDM. The f-DDMwas mounted to the electrical connector and subsequently connected tothe rectifying circuit as shown in left side of Fig. 3c. The f-DDM wassuccessfully operated by flexible wireless power transfer system, andthe optical microscope image of before and after the wireless operationis shown in right side of Fig. 3c. The gold membrane above the mi-croreservoir was dissolved by wireless energy transfer, which indicatesthe stable operation of f-DDM by flexible wireless power transfersystem.

Fig. 3. Flexible wireless power transfer system for the f-DDM. (a) Schematic ofwirelessly powered f-DDM system. (b) The power transfer efficiency result withdifferent antenna to antenna distance (red) and bending radius (blue). (c)Photograph of wirelessly powered f-DDM (left) and optical microscope image ofbefore and after the wirelessly powered operation of f-DDM (right).

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3.4. Analysis of electrochemical reaction of gold membrane

Stable and consistent operation is required for precise control of thedrug release. The release mechanism of the f-DDM is based on thedissolution of the gold sealing membrane under constant DC bias. Thegold membrane provides long-term operation of the f-DDM due to itsexcellent chemical resistance. The drug in the reservoir is delivered bythe dissolution of the gold membrane into biocompatible AuCl4− ions

in the chloride solution (e.g., human body fluid) [23]. This dissolutionprocess is fundamentally governed by the following chemical reactions[39–41]:

+ ⇄ ++ −eAu 4H O Au(H O) 32 2 4

3

+ ⇄ ++ − −Au(H O) 4Cl AuCl 4H O2 4

34 2

First, a linear potential sweep was conducted with respect to the

Fig. 4. Electrochemical analysis of f-DDM operation triggered by applied electric potential. (a) Linear potential sweep results with applied voltage from 0 V to 2 Vwith respect to the Ag/AgCl reference electrode. The high current density peak between 1.0 V and 1.4 V indicates the gold dissolution region. The potential regionabove 1.5 V is dominated by the Cl2 gas evolution reaction. (b) Three-terminal voltage-time curve with constant current density. The constant voltage in thedissolving region shows that the stable gold reaction potential of 1.0 V can be achieved by applying the constant current density 1.1 mA/cm2. (c) Two-terminalvoltage-time curve with the same constant current density demonstrates a stable operation of five cells without the reference electrode. (d) Voltage-time curve fordifferent current density values. The voltage in the dissolving region rises with an increase in current density. Voltage and time deviations as a function of (e) thebending radii (one-way ANOVA, n=4, p=0.799) and (f) bending cycles (one-way ANOVA, n= 4, p= 0.710). The dissolving voltage and end time do not showsignificant degradation.

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reference electrode to accurately observe the potential dependence of theelectrochemical reaction in our f-DDM, as shown in Fig. 4a and b. Fig. 4ashows the sweep results depending on the gold reaction with increasingapplied potential in 1X PBS solution. The potential, which was applied tothe gold sealing membrane serving as a working electrode, was sweptbetween 0V and 2V with respect to the Ag/AgCl reference electrode. Thegold dissolution does not occur in the low-current density region below thepotential of 1.0 V. On the other hand, potential above 1.4 V cannot beutilized for device operation since the Cl2 gas evolution reaction (and O2

evolution if even higher potential is applied) becomes dominant in thisregion. The gold electrodes are effectively dissolved into AuCl4− ions atthe current density peak between 1V and 1.4 V, which is an appropriatepotential range for the gold dissolution in our drug delivery device.Supplementary Fig. 2a depicts the time response of the f-DDM, whichshows gold dissolution without gas evolution at applied potential of 1.0 Vwith respect to the reference electrode. This dependence of the currentdensity as a function of voltage in Fig. 4a is consistent with previousstudies of gold dissolution in chloride solution [42–46].

Fig. 5. In vivo operation of f-DDM implanted beneath the skull in a live mouse. (a) Photograph of the f-DDM implanted live mouse during the brain surgery. (b)Magnified image of the implanted f-DDM. The f-DDM is electrically connected to the adaptor and inserted through a small cranial slit. The Figure insets show aphotograph and a DAPI-labelled fluorescence image of the mouse brain after the removal of the f-DDM. (c) Cross sectional fluorescence images of 50 μm-thick brainslices, and (d) reconstructed three-dimensional image of a mouse brain labelled by two different fluorescent dyes (DiD in blue; DiI in red). Each axis indicatesanterior-posterior (A-P), medial-lateral (M-L), and dorsal-ventral (D-V) directions. (e) The line graph shows the labelled area for each brain slice. The maximumdiffusion area was 0.2 mm2, and the total diffusion volume was 0.058mm3 (DiD) and 0.053mm3 (DiI), respectively.

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Although it can measure the exact potential of the working elec-trodes, the three-terminal method is impractical for actual device op-eration as it is complicated to implement a reference electrode flexiblyembedded in microscale. Instead, a simple and efficient two-terminalmethod was employed for the f-DDM operation without a referenceelectrode. The two-terminal operation using a constant DC voltage bias,however, shows severe deviation in the current density and the dis-solution time as shown in Supplementary Fig. 2b. This non-uniformbehavior of the f-DDM can cause undesired complications in the devicecontrol or even operation failure (see Supplementary Fig. 4a). In con-trast, the two-terminal method using a constant DC current bias is de-termined to be suitable for stable operation of the f-DDM with com-pliance voltage set to avoid undesired high potential. Potential of 1.0 Vwas achieved by applying constant current density of 1.1 mA/cm2,which was previously verified by the three-terminal method, as pre-sented in Fig. 4b. The two-terminal current bias operation for five re-servoirs in Fig. 4c also exhibited highly stable dissolution voltage(1.34 ± 0.017 V) and time (125.9 ± 7.2 s), as shown inSupplementary Fig. 4b. Note that the dissolution voltage is differentfrom the exact potential measured by the reference electrode. Theseoptimized operation conditions provide reliable and precise drug re-lease from the f-DDM. Additionally, Fig. 4d shows the operation currentdensity applied to the f-DDM increased from 1.1mA/cm2 to 5.6 mA/cm2 (i.e. 1 μA to 5 μA per 300× 300 µm2) to adjust the gold dissolutiontime. Based on the electrochemistry data in Fig. 4a and b, the currentdensity of 1.1mA/cm2 was particularly selected as a standard operationcurrent condition that matches the minimal reaction potential of 1.0 V.

For the stable in vivo operation on the curved cerebral cortex ofmice, fatigue tests of the f-DDM in a bending state needs to be carriedout. Reliability tests depending on the bending radius (from 20mm to5mm) was performed by applying constant current density of 1.1 mA/cm2 to examine the mechanical flexibility of the f-DDM. Fig. 4e in-dicates that the dissolution voltage and time were not significantlychanged in the bending states compared to those in the flat state. Theactual strain of the Au film can be calculated as below [16]:

=ε γR

where ε is actual strain and R is bending radius. γ is the distance fromthe mechanical neutral plane, which can be obtained as below:

=+ +

+h

h h h hh h

E E (2 )2(E E )neutral

SU8 SU82

PET PET SU8 PET

SU8 SU8 PET PET

where h is the thickness of each layer and E represents the effectiveYoung's modulus of each material:

=−ν

E E1 2

where ν is the Poisson's ratio and E is the Young's modulus of the ma-terial. The material parameters used for the calculations are (1)ESU8 =3GPa, νSU8 =0.26, hSU8 =30 µm for SU-8 layer and (2)EPET =2GPa, νPET =0.4, hPET =25 µm for PET layer. With bendingradius of 5mm, the actual strain of Au film was calculated as 0.21%,which is much smaller than maximum tensile strain of Au (typically1–2%, depends on strain rate) [47].

Mechanical cycling endurance tests were also performed at abending radius of 5mm. No apparent degradation of the device wasobserved after 1000 bending cycles as shown in Fig. 4f. The statisticalanalysis result in Supplementary Fig. 4 shows that the bending test doeshave significant effect on dissolving voltage and fully dissolved time(one-way ANOVA, n= 4, p=0.799 for voltage and p=0.710 fortime). In addition to its excellent biocompatibility, the f-DDM can showreliable operation on the naturally curved cerebral cortex without anysignificant damage to the brain, because of its outstanding flexibilityand robustness in the bending state.

3.5. Investigation of in vivo drug delivery

In order to examine the performance of the f-DDM in thesubdural space filled with cerebrospinal fluid, in vivo experiments wereconducted with two types of neurotracers; 1,1’-dioctadecyl-3,3,3’,3’-tetra-methylindodicarbocyanine, 4-chlorobenzenesulfonate salt (DiD, λmax, Em

=665nm), and 1,1’-dioctadecyl-3,3,3’,3’-tetramethylindocarbocyanineperchlorate (DiI, λmax, Em =549nm). The neurotracer-loaded f-DDM wasinserted into live mice in their subdural space through a small cranial slit tominimize brain damage from implantation through the removal of thecranium, as shown in Figs. 5a and b [48]. To independently trigger thedissolution of each sealing membrane, each f-DDM reservoir sealing layerwas electrically connected to a customized adaptor as depicted inSupplementary Fig. 5b. The adaptor was attached on the mouse skull usingdental cement, which can effectively prevent infection at least for a shortterm period. The f-DDM was then firmly fixed to the skull of mouse, pro-viding sufficient stability for the targeted administration in freely movingmice. Although the brain may slightly move in a subarachnoid space, the f-DDM can maintain its position on the cortex due to the high mechanicalflexibility.

The mice can freely move after the surgery, and the release ofneurotracers was triggered with electrical stimulation to the separatedreservoir sealing layer in a controlled manner. Consistent with the invitro control experiments, operation current density of 1.1 mA/cm2 wasapplied to minimize damage to the brain. There were no noticeable sideeffects on the general condition of live mice after the administration ofneurotracers. As it was released from each f-DDM reservoir and treatedon the cerebral cortex, the neurotracers were transported to neigh-boring areas via neuronal plasma membranes [49].

Three days after the administration, the labelled brain area wasanalyzed using cross-sectional fluorescent microscopic images and aconsequent reconstructed three-dimensional image. No significantbrain damage was observed over the brain surface as shown in the insetphotograph of Fig. 5b. Furthermore, another inset shows the histolo-gical examination of brain tissues using extra fluorescence image. Asshown in the image of DAPI-stained cell nuclei, there was no criticaldamage to the cerebral cortex after the device implantation and drugadministration. These results are consistent with our previous work,which demonstrated that insertion through a cranial slit can preventbrain damage including structural damage and cell necrosis [16]. Notethat two types of neurotracers from the separated reservoirs were lo-cally delivered into different brain areas, as shown in Fig. 5b. Fig. 5cand d confirm that the brain cells in distinct local cortical areas, whichare less than 1mm apart from each other, were labelled by two dif-ferent neuronal tracing dyes – DiD in blue and DiI in red. Both types ofneurotracers diffused into the cortical area, forming a hemisphericalshape with a diameter of 500 µm. Fig. 5e indicates that the maximumdiffusion length and area were approximately 300 µm and 0.2 mm2,respectively. The total diffusion volume for each neurotracer was cal-culated to be 0.058mm3 for DiD and 0.053mm3 for DiI, respectively. Ingeneral, the drug infusion and diffusion phenomena are strongly de-pendent on the type of drugs. For the development of clinical drugdelivery device, the diffusion properties such as diffusion rate, back-flow, and infusion rate need to be investigated in further research forspecific drug at the pre-clinical trial stage.

In addition to the release of both types of neurotracers, we alsoexamined the performance of a drug-loaded f-DDM. We tested a pos-sible application of f-DDM in the treatment of focal epilepsy with alocalized administration of anti-epileptic drug [50]. To determine itsdiffusional characteristics, in vivo localized drug release of the f-DDMwas demonstrated using florescence-conjugated muscimol (3-hydroxy-5-aminomethylisoxazole), which is effectively used for transmeningealadministration. The muscimol was successfully treated into the targetedbrain area via localized administration, as shown in the fluorescentimage of Supplementary Fig. 7. Furthermore, the therapeutic capabilityof f-DDM was investigated by measuring the EEG signal of epilepsy-

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induced live mice before and after the admnistration of anti-epilepticdrug by f-DDM. Supplementary Fig. 8 shows that the Bicuculine-in-duced spike-and-wave discharge was suppressed after the activation off-DDM, suggesting the clinical applicability of the device. Depending onthe location of targeted region in the brain, the device design can bemodified by several factors that influence diffusional characteristics,such as the size, number, and position of multireservoirs. The localizedadministration through an intracranial therapeutic device is effective,especially for the treatment of focal cortical dysfunctions includingepilepsy, dementia, and ischemia, because systemically-treated centralnervous system (CNS) agents can induce negative neurological effectsin normal brain regions. Since the drug is treated directly onto thebrain, the drug transport is not hindered by blood-brain barrier and thedrug becomes effective with decreased dose [50,51].

In future work, several steps are required to utilize the f-DDM as aclinical method. First, the specification of f-DDM should be adjusted fora specialized treatment. For example, the f-DDM can be integrated withan ECoG device to provide real-time detection of the cortical regiongenerating epileptic activities, which is a target for the drug adminis-tration. The size and the number of reservoirs can be adjusted accordingto the target organs. The high density microreservoir array with morethan 1000 reservoirs in 1×1 cm2 is possible for the long term im-plantation with increased lifetime and drug storage of the device.Second, the specialized forms of drug should be developed for the f-DDM. As described in a previous study for an osteoporosis treatment, ahighly-concentrated neuronal drug that can be effectively loaded in thef-DDM should be developed [19]. Since the brain administration isperformed in the subarachnoid space with cerebrospinal fluid, the drugdelivery method using water-soluble materials can be adopted in the f-DDM system [52]. In addition to traditional drugs, novel biomedicalapproaches such as gene therapy or neuromodulation can be also uti-lized. Finally, for the long term use of the f-DDM, microbattery, flexibleenergy harvester, or wireless power transfer system can be incorporatedas a power source [53–56]. Especially, the integration of demonstratedwireless power transfer system with f-DDM can provide the clinicalsolution for the drug delivery to brain and other soft internal organs.

4. Conclusions

The technology of controlled drug delivery aims to release con-trolled dosage at desired location without affecting nearby organs in aspecific manner. It has advantages like the less side effects, uniformtherapeutic results, and lower dosage compared to oral ingestion andintravenous injection. One of the most important characteristics ofcontrolled drug delivery is the region-specific effect, while traditionaldelivery methods induce systemic effect which may result in undesir-able side effects. Due to the region-specific feature, controlled drugdelivery technology could extend the choice of medication includingthe drugs toxic to normal cells. In case of treatment of epilepsy,transmeningeally delivered Muscimol can prevent the neocortical sei-zures with extremely small dosage of 12.5 mM while it has side effectssuch as hallucination and psychotic symptoms when orally or par-enterally delivered [50].

There has been numerous efforts for the development of controlleddrug delivery device. Santini et al. first demonstrated the prototypemicrochip with precise control of drug release based on the rigid siliconsubstrate [20]. In 2012, Farra et al. reported the first in-human testingof implantable controlled-release microchip with long term operation[19]. Nevertheless, the implantation of the controlled drug deliverydevice to various organs has been complicated because of the bulky sizeand the lack of flexibility. Although there has been a few researches offlexible drug delivery device which is implantable to various bodyparts, these devices are not capable of controlled, timely administration[2,9].

In this report, the controlled drug delivery device imbedded inflexible substrate was developed to overcome these limitations. The

drug release was precisely regulated by the microreservoir structureand the electrochemical dissolution of gold sealing. The operationscheme was optimized through the electrochemical experiment whichis crucial for the timely administration of the drug. The device was firstfabricated on the rigid template and transferred to the flexible substrateusing the laser lift-off technology. The flexible feature of the deviceprovides conformal implantation to the soft organs such as brain, eye,and other internal organs. The implementation of gold sealing mem-brane in the flexible device enables the controlled release operationwith flexible device. Furthermore, the flexible wireless power transfersystem was demonstrated and the f-DDM was successfully operated bynear-field energy transfer expanding the clinical applicability of f-DDM.

The f-DDM can be applied to the treatment of the central nervoussystem (CNS) disorders such as epilepsy, Parkinson's disease, andHuntington's disease. Most of CNS disorders may not be efficientlytreated with oral ingestion or intravenous injection because a highlyselective permeability barrier of the brain called a blood-brain barrier(BBB), complicates drug transport through blood vessels [57]. How-ever, the f-DDM implanted at the brain could deliver the drug to thefocal disease area bypassing the BBB. The localized treatment wouldminimize the side effects of the drug with lower dosage. For example,anticancer drugs attacks not only cancer cells but also normal cellsinducing numerous side effects including vomiting, hair loss, bonemarrow toxicity, and many others. The targeted, localized administra-tion of anticancer drug using f-DDM implanted at the tumor site canminimize the influence to the normal cells result in the reduced sideeffects and also maximized therapeutic effect. The further improvementof flexible drug delivery device and the integration with wireless powersystem and the physiological monitoring sensors would contribute tothe development of precision medicine.

Author contributions

S.H.S., Y.S.K., and D.J.J. contributed equally to this work. S.H.S.,B.H.M., D.J.J., and K.J.L. demonstrated flexible drug delivery device,and investigated its electrical and structural characteristics. Y.S.K andD.K. conducted in vivo drug delivery experiments with flexible drugdelivery device. B.K.Y. analyzed the electrochemical and in vivo ex-periment data. D.H.K., S.K.H., and M.B. analyzed electrical measure-ment result. All authors contributed to discussions and writing of thepaper.

Acknowledgements

This work was supported by the National Research Foundation(NRF) of Korea funded by the Korea government (MSIP) (NRF-2014R1A2A1A12067558 to K. Lee), Wearable Platform MaterialsTechnology Center (WMC) funded by the National Research Foundationof Korea (NRF) Grant of the Korean Government (MSIP) (No.2016R1A5A1009926), and the National Research Foundation (NRF) ofKorea funded by the Ministry of Science, ICT & Future Planning (NRF-2016R1A2B3015167 to D. Kim).

Appendix A. Supporting information

Supplementary data associated with this article can be found in theonline version at http://dx.doi.org/10.1016/j.nanoen.2018.06.015.

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Sang Hyun Sung is a Ph.D. candidate in Materials Scienceand Engineering (MSE) from KAIST, under the supervisionof Prof. Keon Jae Lee. His doctoral research topics includesthe drug delivery device and neuromorphic memory device.

Youngsoo Kim received his B.S. in Korean medicine fromKyung Hee university in 2011, and Ph.D. in neurosciencefrom Korea advanced institute of science and technology(KAIST) in 2016. His current research interest is the clinicalapplication of neuromodulation techniques for the treat-ment of movement disorders.

Daniel J. Joe received a B.S. degree in ElectricalEngineering at the University of Illinois at Urbana-Champaign (UIUC) in 2008 and a Ph.D. degree in Electricaland Computer Engineering at Cornell University in 2014.Currently, he is a BK21 Plus postdoctoral research associatein the Department of Materials Sciences and Engineering atKorea Advanced Institute of Science and Technology(KAIST). His current research interests involve variousflexible thin-film materials and devices including energyharvesters, acoustic sensors, pressure sensors, wirelesspower transfer system, etc.

Beom Ho Mun is a Ph.D., Materior Science andEngineering, 2016, KAIST,Korea. Dissertation: “Fabricationand Characterization of Flexible Memory and Drug DeliveryDevice for Wearable Healthcare System” He is seniormanager of DRAM R&D Process in SK hynix.

Prof. Daesoo Kim received a Ph.D. degree in Genetics andNeuroscience from the Pohang University of Science &Technology (POSTECH) in 1998. He performed post-doctoral research at the State University of New York(SUNY) medical school and worked at Korea Institute ofScience & Technology (KIST) as senior scientist. He is cur-rently studying about the circuit-based mechanism of motorbehavior at Korea Advanced Institute of Science &Technology (KAIST). He has presented as to ‘brain-inspiredsolutions for the crowded earth’ at the World EconomicForum at Davos and won the 3.1 prize from the Samilfoundation.

Prof. Keon Jae Lee received his Ph.D. in Materials Scienceand Engineering (MSE) at University of Illinois, Urbana-Champaign (UIUC). During his Ph.D. at UIUC, he involvedin the first co-invention of “Flexible Single-crystallineInorganic Electronics”, using top-down semiconductors andsoft lithographic transfer. Since 2009, he has been a pro-fessor in MSE at KAIST. His current research topics are self-powered flexible electronic systems including energy har-vesting/storage devices, IoT sensor, LEDs, large scale in-tegration (LSI), high density memory and laser materialinteraction for in-vivo biomedical and flexible application.

S.H. Sung et al. Nano Energy 51 (2018) 102–112

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