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DOI: 10.1002/ ((please add manuscript number)) Article type: Title Microfluidic-release insertion shuttle designed for implanting flexible biomedical electronic devices into live bodies. Author(s), and Corresponding Author(s)* Minjeong Kwon, Kwangsun Song, Dongwuk Jung, Jongho Lee* M.Kwon,K.Song,D.Jung,Prof.J.Lee

School of Mechanical Engineering, Gwangju Institute of Science and Technology (GIST), 123

Cheomdangwagi-ro, Buk-gu, Gwangju 61005, Republic of Korea

E-mail: jong@gist.ac.kr

K.Song,D.Jung,Prof.J.Lee

Research Institute for Solar and Sustainable Energies, Gwangju Institute of Science and

Technology (GIST), 123 Cheomdangwagi-ro, Buk-gu, Gwangju 61005, Republic of Korea

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Abstract

Implantable biomedical devices in flexible forms are attractive as they are more mechanically

compatible with soft live tissues than rigid implants. The flexible implants can bend

comfortably instead of delivering stress or strain to the surrounding live tissues when they are

exposed to external stresses. However, because of the nature of the mechanical properties, the

flexible biomedical implants have one important drawback, difficulty of handling the floppy

implants without bending or wrinkling when inserting through small incision of live tissues.

Here, we present the microfluidic-release insertion shuttle that delivers flexible implants

through the live tissues with added stiffness and releases the implants by dissolving a temporary

adhesive layer with controlled infusing releasing solution through the structures such as post

array and microfluidic channel. Experiments studies with variations in design parameters such

as radii and pitches of the post array provide design guidance of the insertion shuttle. The results

include in vivo experiments with functional flexible electronic implants into a live pig animal

model to demonstrate feasibility of the concept.

1. Introduction

Implantable biomedical devices have been drawing a lot of attention as they can serve real-

time diagnosis and therapy in human bodies. For example, biomedical devices such as

pacemakers[1], cochlear implants[2], gastric stimulators[3], have been already implanted to help

functionalities of live bodies. Furthermore, thin and flexible forms of the implantable devices

such as electrophysiological recording systems[4–6], wound sensors and actuators[7], glucose and

lactate sensors[8,9], blood pressure sensors[10], and drug delivery[11], wireless communication[12]

and energy harvesting systems[13,14] have been also actively developed to alleviate discomfort[15]

to patients by reducing mechanical mismatch between implanted devices and biological

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tissues.[15–18] Although the mechanical compliance of the implanted devices induces minimized

inflammation on live tissues, thus improves long-term stability[14,19–21], the insufficient bending

rigidity can make manipulation of the flexible devices difficult, especially in implantation

procedures, requiring relatively large incision to settle flexible devices.

Recently, many approaches for minimally invasive implantation of flexible devices into bodies

were investigated, particularly for biomedical implants including the flexible neural probes

(FNP)[22] and other silicon electronics[7,23]. One approach is to increase mechanical stiffness

temporarily by coating bioresorbable materials over the flexible devices to insert through an

incision without buckling[24–26]. The devices recover the original flexibility by dissolving the

supporting materials after implantation. Other groups use stiff temporary guides to support

flexible electronic devices while reaching into target organs. The guides include the stiff shank

mounting the sheath type devices[27], the microflap array to switch adhesion on the metal

substrate[28], and the syringe to inject the mesh shape electronics[29]. Although some of these

methods without using bioresorbable materials enable immediate release of devices, deliberate

external manipulation is required to control adhesion or types of implantable devices may be

limited to pass through submillimeter needles. When using bioresorbable materials as

supporting structures, they may need to be thick enough to have sufficient mechanical stiffness

to penetrate through live tissues[30,31] even though it may take much dissolution time to

completely dissolve before starting functioning such as reading biological signals in live bodies.

In many cases, bioresorbable materials such as self-assembled monolayer (SAM) coating[32],

polyethylene glycol (PEG)[33], sacrificial metal[34], and silk fibroin[35] temporarily hold the

flexible devices on the stiff guides while inserting. The procedures require waiting time for the

biomaterials to dissolve to release flexible devices from the temporary stiff guides that are to

be retracted from bodies. It will be very convenient to short the dissolution time for both

surgeons and patients under operations. Here, we present the insertion shuttle that temporarily

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holds a flexible implantable device while inserting and releases the device by dissolving the

adhesive layer with an aid of liquid delivered more efficiently to the interface through the

elastomeric post array and microfluidic channel constituted on the customized stiff metal

substrate. The insertion shuttle we report here has advantages of using a stiff guide for more

convenient handling relatively wide devices without using much degradable materials as

supporting materials, with designs of releasing in short waiting time. We report the designs of

post array structures in different radii and pitches, and evaluated their mechanical adhesion

properties depending on the dissolution time. In vivo implants of the flexible devices such as

the electrocardiogram (ECG) sensor and flexible photovoltaic array into a live animal with the

insertion shuttle reveal the feasibility of the concept.

2. Results and discussion

Figure 1a show an exploded schematic illustration of the microfluidic-release insertion shuttle

designed for implanting a flexible device. The casting and curing processes[36,37]of

polydimethylsiloxane mixture (PDMS, base and curing agent 10:1, Sylgard 184, Dow Corning,

elastic modulus: ~75 MPa) onto SU-8 master molds form the cylindrical post array (height: ~90

µm, diameter: 400 µm, area: 6 mm × 14 mm) in the upper layer (thickness: ~400 µm) and the

microfluidic channel (height: ~90 µm, width:0.2 – 0.6 mm) in the lower layer (thickness: ~ 500

µm). Bonding the two layers with an aid of oxygen plasma treatment connects the channel to

the outlets (diameter: ~1.1 mm, distance: 6 mm) that are made with a biopsy puncher. A thin

film layer (poly-para-xylylene-C, parylene-C, thickness: ~1 µm, elastic modulus: ~4 GPa)

deposited on top of the elastomeric structure reduces tackiness for easier release of a flexible

device.[38,39] The elastomeric structure is mounted on the customized metal support (6 mm × 60

mm, thickness: 0.5 mm). Total thickness of the shuttle (~1.4 mm) could be reduced further to

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minimize lesions in live tissues. More details are in Materials and Methods. A flexible device

is temporarily held on the post array of the integrated insertion shuttle with temporary adhesive

as illustrated in Figure 1b. Releasing solution (Phosphate buffered saline (PBS)) injected

through a tube flows through the channel and outlets, and then spread out through space

between posts (inset), enabling expedite dissolving the adhesive holding the flexible device

temporarily.

Figure 1c shows an optical image of the prepared insertion shuttle holding a flexible device

(solar microcell array (3 × 6), each size: 600 µm × 800 µm, thickness: ~4.1 µm) with an aid of

a temporary adhesive (thickness: ~30 µm) that is a mixture of gelatin, glycerin and water. As

gelatin extracted from animal proteins is known to be biocompatible and bioresorbable [40] ,it is

widely used in drug delivery[41,42], medical glues[43], and hemostatic sponges [44] as well as

coatings for implantable medical electrodes[40,45]. Glycerin also known to be biocompatible[46]is

used as a plasticizer[47]. The weight ratio of gelatin, glycerin, and water was set to be 1:5:5 to

have the mixture in a liquid state over 70 ℃ to form a thin layer by spin-coating on a backside

of a flexible device, and in a gel state at room temperature after contacting the flexible device

onto the post array. Further details are in Materials and Methods. An optical cross sectional

view of the insertion shuttle after mounting a flexible device with an aid of the temporary

adhesive appears in Figure 1d. The flexible device is adhered evenly on the top surface of posts

and the spaces between the posts are secured for infusing releasing solution (PBS) efficiently

as demonstrated in Figure 1e, in which how the releasing solution dyed in red spreads out is

visualized under a transparent film in figure 1e when PBS solution is injected through a tube

(flow rate: 5 µl/s, volume: 20 µl). The injected solution bifurcates at the branch point and

spreads out radially from the outlets. The solution filled around the posts infuses and dissolves

the temporary adhesive layer, resulting in releasing the flexible film mounted on the post array

of the insertion shuttle as demonstrated in Figure 1f. The upper optical image shows the

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insertion shuttle holding the flexible film that is connected to a weight of 20 g over a pulley.

Once the solution is infused, the shuttle releases the film as in the lower image.

The design of the insertion shuttle with the post array and microfluidic channel enables the

implanting process of a flexible device as demonstrated in Figure 2 with a live pig whose skins

are mechanically and anatomically similar to human skins[48–50]. The live pig animal model (age:

9-12 months, weight: 8-12 kg, male) was put in lateral recumbent position with a side skin on

top after injecting a mixture of ketamine (20 mg kg-1, ketamine 50, Yuhan) and zylazine (2 mg

kg-1, Rompun, Yuhan) for general anesthesia. Details of animal preparation are in Materials and

Methods. The implanting process starts with inserting the shuttle (Figure 2a) through the

incision (width: ~7 mm, depth: ~20 mm) made priorly with a sharp surgical blade (No.10,

Paragon) (Figure S1). The flexible device mounted on the shuttle in the image of Figure in 1a

is a flexible electrocardiogram (ECG) device (size: 5 mm × 12 mm, thickness: ~15 µl). Inflow

of the releasing (PBS) solution (flow rate: 5 µl/s, volume: 20 µl, inflowing time: 4 s) through

the microchannel while holding the shuttle down to reduce a normal force helps infusing the

solution between the flexible device and the post array (Figure 2b), lowering adhesion within

about 30-60 seconds. Hereafter, we used the term ‘inflow’ to note filling the empty space

between the posts with the solution during injection and the term ‘infuse’ to note voluntarily

spreading of the releasing solution by dissolving the temporary adhesive between the flexible

film and top surface of the posts during waiting time after stopping injection. Finally, when

retracting the shuttle, we pressed down the inserted shuttle to the body, i.e., away from the skin,

to reduce a normal pressure between the flexible device and shuttle. Lowering normal pressure

results in lower friction (Figure S3 in Supporting Information) between the flexible device and

shuttle than the friction (1.65 – 4.57 mN/mm2) [28]between the flexible device and skin, thus,

leaving the flexible device under the skin (Figure 2c).

The adhesive and release properties of the insertion shuttle with various post array designs are

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characterized with the custom lap shear measurement setup as in Figure 3a. The motorized 4-

axis (X, Y, Z, and rotation) microstage pulls the insertion shuttle (glass support, speed: 100

µm/s, blue arrow) on which the post array is adhered to the flexible PI film (thickness: 12.5 µm,

length: 30 mm, and width: 5 mm) with an aid of the gelatin-glycerin adhesive (thickness: ~30

µm). The other end of the flexible film is fixed to the force sensor, enabling measurements of

the lap shear failure forces (FL, red arrow) between the postal array and the flexible PI film.

Further details of the setup are in Materials and Methods. Various post array designs (from (i)

to (vii)) with different radii (0.2-0.4 mm) and pitches (0.6-1.2 mm) used in the measurements

appear in Figure 3b. For each design, a lateral distribution of post, top view image, and areal

ratio are tabulated in each cell. An areal ratio is defined as a ratio of the total top surface area

of posts (AT) to the device area (AD) as illustrated in the second cell of Figure 3b.

Figure 3c shows the measurement results (n = 5) of lap shear failure forces (FL) of the designs

(i: red triangle, ii: red square, iii: green square, iv: blue square, v: red circle, vi: green circle, vii:

blue circle) without injecting the releasing(PBS) solution. The lap shear failure forces (4.0 –

17.8 mN/mm2) are irrelevant to the radii or pitches but they are rather linearly depending on

areal ratios (AT/AD = 9.2 – 51.1%) as no complex mechanical deformation such as bending is

expected for the posts whose aspect ratios are lower than 0.3.[51] The height of the posts (~90

µm) is high enough to secure spaces for releasing solution. Proper post array designs are

determined to have higher adhesion than the friction between the film and live tissue (1.65 –

4.57 mN/mm2, grayed in Figure 3c[28]) to hold a flexible device while inserting. After inflowing

(injection starts at -4s and stops at 0s in Figure 3d) the releasing (PBS) solution, the lap shear

failure forces decrease as waiting time (infusing time) becomes longer as shown in Figure 3d.

For example, after waiting for 60s to allow infusing the solution, the lap shear failure force (FL

= 0.6 mN/mm2 after 60s for the design (i)) between the PI film and shuttle becomes lower than

the friction force (1.65 – 4.57 mN/mm2) between the PI film and tissue, leaving the PI film with

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the tissue while retracting the shuttle. The rate of decay, however, for a given period of time,

relies on that how small the contact surface (sum of the top surface of posts) is separated. For

example, the designs (i: red triangle, radius:0.2 mm, pitch: 0.6 mm) and (vii: blue circle, radius:

0.4 mm, pitch: 1.2 mm) have similar initial lap shear failure forces (11.0 mN/mm2 vs. 12.1

mN/mm2) as areal ratios are similar (34.0% vs. 36.2%) but the lap shear failure force of the

design (i) decreases more rapidly(11.0 mN/mm2 à 0.9 mN/mm2) than the design (vii)

does(12.1 mN/mm2 à 4.2 mN/mm2) for the same period of waiting time (30 s) because the

smaller radius requires less infusing distance of the releasing solution, which is supported by

that the normalized lap shear failure forces can be fitted with the model for remaining adhesives

immersed in dissolution solution. The time constants are short for the smaller radii. More details

are in Supporting Information (Figure S2 and Table S1).

Infusing of the releasing solution is captured in the time-lapse optical images (scale bars: 2

mm) of a flexible PI film mounted onto the post array (radius: 0.4 mm, pitch: 1.2 mm) in Figure

3e. After stopping injecting the PBS solution, the PI film are embossed (as shown at 5 s, 10 s,

and 30s) around the posts during infusing as the solution starts to wet and dissolve the gelatin-

glycerin adhesive from the circular top edges of the posts. After fully dissolving the adhesive

through the top surfaces of the posts, the embossings disappear at 60 s. The dissolved adhesive

attracts water (PBS is a water-based solution) as demonstrated with the spreading contact angles

(figure 3f) quantified from the time-lapse optical images (n = 6) of water droplets (deionized

water, 25 ℃, volume: ~ 8 µl) on the gelatin-glycerin adhesive layer applied on a PI film. The

contact angle (~ 65° at 0s) reduces down to 31° at 60 s as the droplet dissolves the adhesive

layer.

The structured microfluidic-release insertion shuttle reduces adhesion with a higher ratio as

the PBS solution spreads more efficiently, compared to the non-structured insertion shuttle

simply covered with the same material (parylene-C, thickness: ~1 µm) over smooth PDMS

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slabs (size: 5 mm × 12 mm × 0.4 mm) without any surface structures (Figure S4). The

microfluidic-release insertion shuttle can deliver flexible medical electronics subcutaneously as

demonstrated in a live animal model. Figure 4 shows a ventral view of locations where the

flexible devices are inserted in a pig model. The flexible electrocardiogram (ECG) sensor[28]

(size: 5 mm × 12 mm, thickness: ~14.5 µm) consisting of electrodes (size: 1.5 mm × 1.8 mm,

period: 3.5 mm) and traces (width: 0.5 mm, pitch: 0.5 mm) was inserted, facing toward the

heart, on the left lateral pectoral region, near to the heart, along the anterior axillary line of a

live pig model (age: 9-12 months, weight: 8-12 kg, male) under general anesthesia as shown

with the upper image of Figure 4b. More details of the ECG sensors are in Materials and

Methods. For comparison, the identical ECG was attached on the shaved skin in the similar

location (lower image of Figure 4b) with highly conductive electrode gel (Signagel, Parker

Laboratories). The ECG signals (black line in Figure 4c) from the inserted device have

distinguishable (~67 bpm) typical forms of QRS complex and inverse T wave while the attached

sensor reads indistinguishable signals (red line) possibly due to crosstalk or low signal

amplitude on the skin.

As another application, we used the microfluidic-release insertion shuttle to implant a flexible

photovoltaic (IPV) device[52] on the right hypochondrium. The flexible IPV device is bendable

on a cylinder of a radius 4 mm (Figure 4d), and mechanically durable to maintain the conversion

efficiency and power for repetitive bending and unbending upto 1,000 cycles on the same

cylinder (radius: 4 mm) as in Figure 4e. The flexible IPV device implanted with the insertion

shuttle (Figure 4f) still generates electrical power (0.35 mW, short-circuit current density: 2.3

mA/cm2, conversion efficiency: 4.6%) under the skin (red line in Figure 4g) by capturing light

penetrated through the skin although the electrical performances are lower than the

measurements under standard AM1.5G illumination (100 mW/cm2, LCS-100, Oriel

Instruments) out of the skin (black line in Figure 4g).

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3. Conclusion

The insertion shuttle reported here presents an effective strategy to control adhesion for

expedite release of flexible medical devices in live bodies when implanting the flexible devices

through a small incision. Studies on the mechanical properties based on experimental results

with various radii and pitches of the post arrays provide important guidance for designing the

insertion shuttle. Demonstrations with the flexible ECG sensor and flexible photovoltaic

devices prove the capability of the injector for functioning devices. Further studies and designs

of the structures along with development of improved biocompatible adhesives should lead to

practical uses of the concept for various kinds of flexible medical implants whose

functionalities and capabilities have been actively expanding these days.

4. Materials and Methods

Preparation of the microfluidic-release insertion shuttle: The fabrication of the microfluidic-

release insertion shuttle began with preparing SU-8 master molds to replicate a microfluidic

channels and post structures. A thin SU-8 epoxy layer (thickness: ~20 µm, Microchem) was

spin-coated on a transparent glass substrate (Micro glass substrate, Matsunami Glass) to

improve adhesion to a thick SU-8 layer (thickness: ~90 µm) in which the microfluidic channels

were formed by photolithography. The separate SU-8 molds for post structures were prepared

in the same manner. The elastomeric microfluidic channel layer (thickness: ~500 µm) and the

post array layer (thickness: ~400 µm) were formed by spin-casting mixed silicone prepolymer

and curing agent (PDMS, 10:1, Sylgard 184, Dow corning) on the molds, curing at 85°C for an

hour, and separating from the molds. The elastomeric mold layers were transferred onto

temporary supporting films (Polyethylene terephthalate, thickness: 100 µm). After two outlet

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holes (diameter: ~1.1 mm) were made with a biopsy punch (BPP-10F, Kai medical) on post

array layer, the microfluidic channels layer and the post array layer were bonded with an aid of

oxygen plasma treatment, followed by connecting a tube (OD: 0.6 mm, ID: 0.4 mm, PFA, AS

ONE) to the microfluidic channel. Finally, the thin film layer (poly-para-xylylene-C, parylene-

C, thickness: ~1 µm) was deposited on top of the elastomeric structure mounted on the stainless

support (size: 6 mm × 60 mm, thickness: 0.5 mm, and fillet radius: 2 mm) customized with

wire-cutting machining.

Mounting a flexible device on the shuttle: The temporary adhesive to hold a flexible device on

the shuttle was prepared by stirring gelatin powder (Type B, Sigma), glycerin (BioXtra, Sigma-

Aldrich), and deionized water with a weight ratio of 1:5:5 on a hot plate at 70 °C. A flexible

device whose backside spin-coated (thickness: ~30 µm) with the dispersed temporary adhesive

was mounted on the post array of the insertion shuttle. The flexible device was pressed with a

glass slide (size: 18 mm × 25 mm, weight: ~1.2 g) for uniform contact with the post array for

20 minutes of resting time at room temperature.

Preparation of the animal models: We conducted the experiments with the pig animal models

(age: 9-12 months, weight: 8-12 kg, Medi Kinetics) because of their similarity to human skins

in mechanically and anatomically.[48–50] All the experiments on the animal models were

conducted in accordance with GIST IACUC (GIST-2017-075) approval. The animal tests start

with intramuscular injection of mixture of ketamine (20 mg kg-1, Ketamine 50, Yuhan) and

zylazine (2 mg kg-1, Rompun, Yuhan) for general anesthesia. With the animal model placed in

a lateral recumbent position, its skin was sterilized with povidone iodine (Povidin, Firson)

ahead of making incision and subcutaneous insertion. The ophthalmic ointment (Liposic,

Bausch & Lomb) was applied to the pigs’ eyes regularly to keep long lasting hydration of the

cornea to avoid irritation. An incision (size: 7 mm × 20 mm, depth: ~2 mm) on the previously

marked incision (~10 mm) was made with separate surgical blade (No.10, Paragon) before

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inserting a flexible device subcutaneously. Too much liquid in the incision in cases of delaying

too much time (> 1 min) or repetitive trials into one incision may result in unsuccessful releasing

a flexible device in the body because the incised skin becomes too much slippery. Breathing

motion of the animal model may also affect implantation process adversely when inserting a

device around bones, e.g., chest bones, because too much normal pressure can be applied while

retracting the shuttle.

Measurements of the mechanical characteristics: The lap shear failure forces of the flexible

device mounted on the post array with an aid of the temporary adhesives were evaluated for the

various designs of the post arrays using a force sensor (range: ±150 g, Transducer techniques,

USA) while moving the 4-axis (X, Y, Z, and rotation) motorized microstage laterally (100

µm/sec). The force sensor connected to the flexible device is mounted on a 3-axis manual

microstage that aligns the flexible device to the insertion shuttle fixed on the motorized

microstage. We used polyimide (PI) films (5 mm × 30 mm, thickness: 12.5 µm) as dummy

flexible devices for characterizing the lap shear failure forces. To see the effects of the PBS

infusion, a syringe pump (Pump 11 elite infusion/withdrawal syringe pump, Harvard Apparatus)

introduced the PBS through the tube.

Preparation of the flexible implantable devices: The flexible electrocardiographic (ECG) (size:

5 mm × 12 mm) sensors were prepared by depositing and patterning metals (Ti: 30 nm/ Au:

150 nm) on the PI film (thickness: 12.5 µm) to form the electrodes (size: 1.5 mm × 1.8 mm,

period: 3.5 mm) and traces (width: 5 mm), then encapsulated with a SU-8 thin layer (~2 µm)

except the electrodes to be opened, as reported previously[28]. The traces are connected through

the anisotropic conductive film (pitch: 5 mm) to the oscilloscope (DSO-X 2024A, Keysight

technologies) via differential amplifier (PSL-iECG2, Physiolab) for monitoring and acquiring

ECG signals.

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As an implantable power generator, arrays of the solar microcells (size: 800 µm × 600 µm,

thickness: ~4.1 µm) were fabricated from epitaxially grown on GaAs wafers, and transferred

onto a flexible substrates (thickness: 12.5 µm) with an aid of a spin-casted thin SU-8 adhesive

layer (thickness: ~2 µm) as reported previously[53]. After interconnecting the solar microcells

in series (×3) and parallel (×6) by defining additional metal layers, the array was encapsulated

with a transparent SU-8 layer (thickness: 6 µm) with openings to expose metal electrodes (Ti:

30 nm/ Au: 200 nm). The electrical performance (current-voltage characteristic) of the

implantable solar microcells array was measured with a source measurement unit (B2902A,

Keysight technologies), under the standard AM 1.5G illumination (100 mW/cm2, LCS-100,

ORIEL Instruments).

Supporting Information Supporting Information is available from the Wiley Online Library or from the author.

Acknowledgements

M.K and K.S contributed equally to this work. This work was supported by the National

Research Foundation of Korea (NRF) grant funded by the Korea government (No.

2016R1A2B4012854), the Korea Institute of Energy Technology Evaluation and Planning

(KETEP) funded by the MOTIE (No. 20163030013380), and by the GIST Research Institute

(GRI) project through a grant provided by GIST.

Received: ((will be filled in by the editorial staff)) Revised: ((will be filled in by the editorial staff))

Published online: ((will be filled in by the editorial staff))

References

[1] D. A. Nathan, S. Center, C. Wu, W. Keller, Am. J. Cardiol. 1963, 11, 362.

[2] B. J. Gantz, C. Turner, K. E. Gfeller, M. W. Lowder, Laryngoscope 2005, 115, 796.

14

[3] T. L. Abell, E. Van Cutsem, H. Abrahamsson, J. D. Huizinga, J. W. Konturek, J. P.

Galmiche, G. Voeller, L. Filez, B. Everts, W. E. Waterfall, W. Domschke, S. B. Des

Varannes, B. O. Familoni, I. M. Bourgeois, J. Janssens, G. Tougas, Digestion 2002, 66,

204.

[4] H. Fang, K. J. Yu, C. Gloschat, Z. Yang, E. Song, C. H. Chiang, J. Zhao, S. M. Won,

S. Xu, M. Trumpis, Y. Zhong, S. W. Han, Y. Xue, D. Xu, S. W. Choi, G.

Cauwenberghs, M. Kay, Y. Huang, J. Viventi, I. R. Efimov, J. A. Rogers, Nat. Biomed.

Eng. 2017, 1, DOI 10.1038/s41551-017-0038.

[5] K. J. Yu, D. Kuzum, S. W. Hwang, B. H. Kim, H. Juul, N. H. Kim, S. M. Won, K.

Chiang, M. Trumpis, A. G. Richardson, H. Cheng, H. Fang, M. Thompson, H. Bink, D.

Talos, K. J. Seo, H. N. Lee, S. K. Kang, J. H. Kim, J. Y. Lee, Y. Huang, F. E. Jensen,

M. A. Dichter, T. H. Lucas, J. Viventi, B. Litt, J. A. Rogers, Nat. Mater. 2016, 15, 782.

[6] P. J., C. S., J. A.H., L. S.-Y., R. S., S. J., S. K., Y. S., L. C., K. K.-W., C. H.R., K. S.J.,

S. P., H. W., J. S., L. H.-J., L. N., C. S.H., S. M., L. N., J. M.E., H. T., K. D.-H., Sci.

Transl. Med. 2016, 8, 344ra86.

[7] D. H. Kim, S. Wang, H. Keum, R. Ghaffari, Y. S. Kim, H. Tao, B. Panilaitis, M. Li, Z.

Kang, F. Omenetto, Y. Huang, J. A. Rogers, Small 2012, 8, 3263.

[8] A. Guiseppi-Elie, S. Brahim, G. Slaughter, K. R. Ward, IEEE Sens. J. 2005, 5, 345.

[9] K. W. Johnson, J. J. Mastrototaro, D. C. Howey, R. L. Brunelle, P. L. Burden-Brady,

N. A. Bryan, C. C. Andrew, H. M. Rowe, D. J. Allen, B. W. Noffke, W. C. McMahan,

R. J. Morff, D. Lipson, R. S. Nevin, Biosens. Bioelectron. 1992, 7, 709.

[10] M. Theodor, D. Ruh, J. Fiala, K. Förster, C. Heilmann, Y. Manoli, F. Beyersdorf, H.

Zappe, A. Seifert, Biomed. Microdevices 2013, 15, 811.

15

[11] D. R. Kipke, D. S. Pellinen, R. J. Vetter, 2002 IEEE Int. Symp. Circuits Syst. Proc.

(Cat. No.02CH37353) n.d., 4, IV.

[12] G. Shin, A. M. Gomez, R. Al-hasani, J. S. Ha, M. R. Bruchas, J. A. Rogers, 2017, 509.

[13] K. Song, J. H. Han, H. C. Yang, K. Il Nam, J. Lee, Biosens. Bioelectron. 2017, 92, 364.

[14] L. Lu, Z. Yang, K. Meacham, C. Cvetkovic, E. A. Corbin, A. Vázquez-Guardado, M.

Xue, L. Yin, J. Boroumand, G. Pakeltis, T. Sang, K. J. Yu, D. Chanda, R. Bashir, R.

W. Gereau, X. Sheng, J. A. Rogers, Adv. Energy Mater. 2018, 8, 1.

[15] K. Bazaka, M. Jacob, Implantable Devices: Issues and Challenges, 2012.

[16] I. R. Minev, P. Musienko, A. Hirsch, Q. Barraud, N. Wenger, E. M. Moraud, J.

Gandar, M. Capogrosso, T. Milekovic, L. Asboth, R. F. Torres, N. Vachicouras, Q.

Liu, N. Pavlova, S. Duis, A. Larmagnac, J. Vörös, S. Micera, Z. Suo, G. Courtine, S. P.

Lacour, Science (80-. ). 2015, 347, 159.

[17] A. Weltman, J. Yoo, E. Meng, Micromachines 2016, 7, DOI 10.3390/mi7100180.

[18] J. Rivnay, H. Wang, L. Fenno, K. Deisseroth, G. G. Malliaras, Sci. Adv. 2017, 3, DOI

10.1126/sciadv.1601649.

[19] S. Takeuchi, D. Ziegler, Y. Yoshida, K. Mabuchi, T. Suzuki, Lab Chip 2005, 5, 519.

[20] J. Shi, Y. Fang, Adv. Mater. 2018, 1804895, 1804895.

[21] H. Wang, X. Ma, Y. Hao, Adv. Mater. Interfaces 2017, 4, 1.

[22] J. A. Rogers, R. Ghaffari, D.-H. Kim, Microsystems and Nanosystems Stretchable

Bioelectronics for Medical Devices and Systems, 2016.

[23] D. H. Kim, Y. S. Kim, J. Amsden, B. Panilaitis, D. L. Kaplan, F. G. Omenetto, M. R.

Zakin, J. A. Rogers, Appl. Phys. Lett. 2009, 95, 1.

16

[24] T. Suzuki, K. Mabuchi, S. Takeuchi, Int. IEEE/EMBS Conf. Neural Eng. NER 2003,

2003–Janua, 154.

[25] Z. Xiang, S. C. Yen, N. Xue, T. Sun, W. M. Tsang, S. Zhang, L. De Liao, N. V.

Thakor, C. Lee, J. Micromechanics Microengineering 2014, 24, DOI 10.1088/0960-

1317/24/6/065015.

[26] D. Lewitus, K. L. Smith, W. Shain, J. Kohn, Acta Biomater. 2011, 7, 2483.

[27] B. J. Kim, J. T. W. Kuo, S. A. Hara, C. D. Lee, L. Yu, C. A. Gutierrez, T. Q. Hoang, V.

Pikov, E. Meng, J. Neural Eng. 2013, 10, DOI 10.1088/1741-2560/10/4/045002.

[28] K. Song, J. Kim, S. Cho, N. Kim, D. Jung, H. Choo, J. Lee, Adv. Healthc. Mater. 2018,

7, 1.

[29] J. Liu, T. M. Fu, Z. Cheng, G. Hong, T. Zhou, L. Jin, M. Duvvuri, Z. Jiang, P. Kruskal,

C. Xie, Z. Suo, Y. Fang, C. M. Lieber, Nat. Nanotechnol. 2015, 10, 629.

[30] M. Pawlaczyk, M. Lelonkiewicz, M. Wieczorowski, 2013, 302.

[31] S. Budday, R. Nay, R. De Rooij, P. Steinmann, T. Wyrobek, C. Ovaert, E. Kuhl, E.

Prairie, 2016, 318.

[32] T. D. Y. Kozai, D. R. Kipke, J. Neurosci. Methods 2009, 184, 199.

[33] S. Felix, K. Shah, D. George, V. Tolosa, A. Tooker, T. Delima, S. Pannu, 2012, 871.

[34] X. Jiao, Y. Wang, Q. Qing, Nano Lett. 2017, 17, 7315.

[35] J. Wyman, J. P. Changeux, D. Filmer, T. M. Jovin, W. Baehr, J. J. Holbrook, N.

Dattagupta, D. M. Crothers, G. W. Hatfield, R. Bruinsma, T. Maniatis, S. C. Harrison,

A. J. Spakowitz, P. C. Blainey, C. M. Schroeder, X. S. Xie, T. Strzelecka, L. F. Dorner,

I. Schildkraut, A. K. Aggarwal, S. Bailey, T. A. Steitz, L. Finzi, C. Bustamante, C. T.

17

Martin, S. S. Patel, A. Kumar, S. S. Patel, S. Oehler, A. K. Aggarwal, S. Stayrook, A.

Rosenberg, M. Lewis, J. Widom, J. Hynes, A. Szabo, C. Bustamante, 2013, 819.

[36] A. Folch, A. Ayon, O. Hurtado, M. a Schmidt, M. Toner, J. Biomech. Eng. 1999, 121,

28.

[37] H. E. Jeong, P. Kim, M. K. Kwak, C. H. Seo, K. Y. Suh, Small 2007, 3, 778.

[38] T. Y. Chang, V. G. Yadav, S. De Leo, A. Mohedas, B. Rajalingam, C. L. Chen, S.

Selvarasah, M. R. Dokmeci, A. Khademhosseini, Langmuir 2007, 23, 11718.

[39] H. Y. Sun, C. Y. Fang, T. J. Lin, Y. C. Chen, C. Y. Lin, H. Y. Ho, M. H. C. Chen, J.

Yu, D. J. Lee, C. H. Chang, H. Y. Chen, Adv. Mater. Interfaces 2014, 1, 1.

[40] J. A. Witteveen, D. B. Suyatin, L. Gällentoft, J. Schouenborg, N. Danielsen, C. N.

Prinz, J. Vac. Sci. Technol. B, Nanotechnol. Microelectron. Mater. Process. Meas.

Phenom. 2010, 28, C6K13.

[41] S. Schrade, K. Zeyer, M. Keck, M. Keller, M. Zähner, J. Mohn, Soft Gelatin Capsules

(Softgels), 2018.

[42] M. Foox, M. Zilberman, Expert Opin. Drug Deliv. 2015, 12, 1547.

[43] Y. Otani, Y. Tabata, Y. Ikada, Biomaterials 1996, 17, 1387.

[44] E. Cenni, G. Ciapetti, S. Stea, A. Corradini, F. Carozzi, J. Biomater. Sci. Polym. Ed.

2000, 11, 685.

[45] J. Agorelius, F. Tsanakalis, A. Friberg, P. T. Thorbergsson, L. M. E. Pettersson, J.

Schouenborg, Front. Neurosci. 2015, 9, 1.

[46] Z. Mao, X. Wen, J. Wang, D. Zhu, 2009, 7519, 75191N.

[47] P. Bergo, R. A. Carvalho, A. C. S. Vadala, V. C. I. Guevara, P. J. A. Sobral, 2010, 753.

18

[48] A. Summerfield, F. Meurens, M. E. Ricklin, Mol. Immunol. 2015, 66, 14.

[49] O. A. Shergold, N. A. Fleck, D. Radford, Int. J. Impact Eng. 2006, 32, 1384.

[50] J. Lim, J. Hong, W. W. Chen, T. Weerasooriya, Int. J. Impact Eng. 2011, 38, 130.

[51] R. K. Kramer, C. Majidi, R. J. Wood, Adv. Mater. 2010, 22, 3700.

[52] K. Song, J. H. Han, T. Lim, N. Kim, S. Shin, J. Kim, H. Choo, S. Jeong, Y. Kim, Z. L.

Wang, J. Lee, 2016, 1572.

[53] J. Kim, N. Kim, M. Kwon, J. Lee, ACS Appl. Mater. Interfaces 2017, 9, 25700.

[54] K. Kucera, A. Jessop, N. Alvarez, D. Gortler, D. Light, G. Washington, I. S. Nw, 2018,

71, 63.

[55] A. K. KH Ramteke, PA Dighe, SV Patil - Sch. Acad. J. Pharm 2014, 3, 388.

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Figure 1. Schematic and optical images of the microfluidic-release insertion shuttle. a)

Exploded and b) integrated illustrations of the insertion shuttle. c) An optical image of the

integrated insertion shuttle holding a flexible device (implantable photovoltaic (IPV) device)

with an aid of a temporary adhesive (A mixture of gelatin and glycerin). Inset scale bar: 2 mm.

d) A cross-sectional view of the insertion shuttle holding the flexible device. Only the top

surfaces of the posts contact and hold the flexible device with the adhesive. PBS solution can

flow through between the posts (height: ~90 µm). e) Sequential top view images of the insertion

shuttle through a dummy transparent flexible device (PI film) while infusing PBS solution dyed

in red to show how the PBS solution spreads out. f) (upper) Optical image of the insertion

shuttle holding the flexible PI film that is connected through a string to the hanging weight (20

g) over a pulley. (lower)The shuttle dismounts the flexible film as the inflow of solution (dyed

in red) spreads out and infuses between the posts and flexible device.

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Figure 2. Schematic illustrations and optical images of the process of implanting a flexible

device under the skin of a live pig model with the insertion shuttle. a) The shuttle on which the

flexible device is adhered is put through the subcutaneous incision (size: 7 mm × 20 mm, depth:

~2 mm) made with a surgical blade priorly. b) PBS solution flows in the microchannel with a

syringe pump (flow rate: 5 µl/sec, volume: 20 µl). After waiting about 30-60 seconds, c) the

insertion shuttle is retracted, leaving the flexible device under the skin.

21

Figure 3. Characteristics of the insertion shuttle. a) Experimental setup for characterizing lap

shear failure forces of the insertion shuttle that has various post array designs. b) Various post

array designs, optical images and areal ratios (=AT/AD) in different radii (0.2-0.4 mm) and

pitches (0.6-1.2 mm). Scale bar: 2 mm. The areal ratio is a ratio of the total top surface area of

the contacting posts (AT, colored solid circles) over the flexible device area (AD, dashed

rectangle). c) Measurement results of the lap shear failure force FL (n = 5) with respect to the

areal ratios determined by the various post array designs. d) Measurement results of FL

depending on waiting times (0, 30, 60, and 120 sec) after inflow (flow rate: 5 µl/sec, volume:

20 µl) of the PBS solution. FL at -4 sec is the force without inflowing the PBS solution. e)

Magnified view images of a flexible PI film initially adhered to the shuttle (Post radius: 0.4 mm,

pitch: 1.2 mm) (scale bars: 2 mm). While the PBS solution infuses, the PI film around the posts

are embossed (5, 10, and 30 s). After fully infusing, the flexible device appears to be slightly

floating (60 s). f) Measurements results of contact angles of water drops on the gelatin-glycerin

adhesive layer on the PI film (n= 5). Images: time-lapse contact angles (scale bar: 1 mm).

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Figure 4. Demonstrations of inserting functioning flexible devices subcutaneously with the

microfluidic-release insertion shuttle in a live pig model. a) The locations of the flexible devices

being implanted in a pig model. b) Optical images of the flexible ECG sensor being inserted

under the skin (upper) and attached on the skin (lower). c) Signals measured with the ECG

sensor inserted under the skin (black) and attached on the skin (red). The signal obtained from

the inserted ECG sensor captures typical shapes of pulses. d) The flexible IPV device bent on

a cylinder of a radius of 4 mm. e) Conversion efficiency and electrical power of the flexible

IPV after cycles of bending and unbending on the cylinder (R= 4 mm). f) Optical image of

subcutaneously inserting the flexible IPV device in a live pig. g) Current-voltage curves of the

flexible IPV device measured (black) out of the skin and (red) under the skin.