Inkjet-Printed Neural Electrodes with Mechanically Gradient Structure

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Cite This: ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

Inkjet-Printed Neural Electrodes with Mechanically Gradient Structure Nana Kokubo,† Masashi Arake,⊥ Kento Yamagishi,‡ Yuji Morimoto,¶ Shinji Takeoka,† Hiroyuki Ohta,*,¶ and Toshinori Fujie*,§,#,∥

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Department of Life Science and Medical Bioscience, Graduate School of Advanced Science and Engineering, ‡Research Organization for Nano & Life Innovation, and §Waseda Institute for Advanced Study, Waseda University, 1-104, Totsuka-cho, Shinjuku, Tokyo 169-8050, Japan ⊥ Department of Integrative Physiology and Bio-Nano Medicine and ¶Department of Physiology, National Defense Medical College, 3-2 Namiki, Tokorozawa, Saitama 359-8513, Japan # PRESTO, Japan Science and Technology Agency, 4-1-8, Honcho, Kawaguchi-shi, Saitama 332-0012, Japan ∥ School of Life Science and Technology, Tokyo Institute of Technology, B-50, 4259 Nagatsuta-cho, Midori-ku, Yokohama 226-8501, Japan S Supporting Information *

ABSTRACT: Flexible materials are important for the development of neural probes in recording stable signals (spikes) in vivo. Here, we present inkjet-printed, flexible neural probes for spike recording by using polymeric thin films. The neural probes were constructed from 400 nm-thick poly(D,L-lactic acid) nanofilms, inkjet-printed lines consisting of Au and poly(3,4-ethylenedioxythiophene):polystyrenesulfonate nanoinks, and fluoropolymer layers. Microelectrodes were exposed by cutting the edge with a razor. The 6 μm-thick probes were connected to the external amplifiers by gradual increase of stiffness with thickness-dependent manner. The probe was formed into a needle shape, which recorded spikes from mouse thalamus in vivo. KEYWORDS: polymer nanofilm, inkjet printing, printed electronics, neural electrode, optogenetics

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electronics made of soft materials (e.g., polyimide and Parylene C [Young’s modulus of 3.2 GPa]4 and SU-8 [2 GPa]).6 The flexibility of electronics based on polymer thin films makes them suitable for coupling with neuronal cells in fragile brain tissues, and these electronics have been used to measure not only electrocorticogram signals and local field potentials (LFPs)7 but also spikes from neurons.8,9 However, such flexibility causes poor strength in operation, making it difficult to insert the probe into rodent brains, and also causes mechanical failure between the microfabricated probe and rigid external devices. Because of these issues, shuttle devices such as a needle, tube, and syringe are often used to guide the electronics into the desired tissue regions.8,9 In addition, previous reports have proposed cumbersome procedures to connect flexible electronics to the rigid external amplifier by using an anisotropic conductive film with vacuum and high temperature or heating processing.8 Therefore, flexible

iomedical materials including smart engineering technologies are required for the development of implantable electronic devices to minimize invasiveness, economic constraints, and manufacturing problems.1 Neural probes inserted into deep brain tissues are important tools for electrophysiological recording of neuronal signals to improve our collective understanding of brain functions in neurological disorders.2 However, conventional neural probes such as silicon probes have several drawbacks including high rigidity compared with soft brain tissues and poor mechanical properties during manipulation. When using silicon probes, the operator must handle the probe carefully because its tip is extremely small, sharp, and brittle, so it can be easily damaged if touched by a finger or dropped on the floor. Moreover, the mechanical mismatch between rigid conventional electrodes (e.g., Young’s modulus for silicon probes: ∼150 GPa)3 and soft brain tissues (∼100 kPa)4 causes tissue damage from micromotion of the implanted electrodes.4,5 Therefore, alternative materials and engineering technologies to fabricate neural electrodes have been investigated.5 Recent reports have shown alternatives to silicon and microwire electronics in the form of polymer-based flexible © XXXX American Chemical Society

Received: September 30, 2018 Accepted: December 7, 2018 Published: December 7, 2018 A

DOI: 10.1021/acsabm.8b00574 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

Letter

ACS Applied Bio Materials

Figure 1. Overall view of the flexible polymer sheet electrode. (a) Picture of a flexible sheet electrode. (b) Schematics of the layered structure of the sheet electrode and the cross-sectional view. Blue, black, sky blue, yellow, gray, and dark blue represent the PDLLA nanofilm, conductive lines (Au and PEDOT: PSS), CYTOP, polyimide film (12.5 μm), double-sided tape, and precut FFC, respectively. (i, ii, iii) Schematics in the black dashed box and (iv) in the red dashed box indicate the cross-sections of the black and red lines in panel b, which represents the electrode part insulated with CYTOP, polyimide film-attached printed nanofilm between the electrode and the connection part, and the connection part, and overall structure, respectively. (c) Fabrication of a sheet electrode. Au and PEDOT:PSS lines were inkjet-printed on a PDLLA nanofilm. CYTOP was dipcoated on both sides of the printed PDLLA nanofilm. The trimmed CYTOP-coated PDLLA nanofilm was 3 mm wide and 50 mm long. The thicknesses of the electrode part (i) and the connection part (ii) were ∼6 μm and ∼13 μm, respectively.

also integrated into an optical fiber (i.e., optrode) to enable optogenetic recording. (iv) To connect the sheet electrode to the external amplifiers, the stiffness was gradually increased from the flexible nanofilm to the rigid device and fabricated preassembled connection. The fabricated needle electrode or optrode successfully recorded spontaneously and optogenetically evoked individual neuron spikes in rodents. To the best of our knowledge, this is the first report to fabricate a spikerecording electrode by inkjet-printing process. We demonstrate how to solve the connection problem between thin/small neural electrodes and external devices by gradually increasing the mechanical stiffness of the neural probe with stepwise increment of the layer thickness from the electrode (micrometer scale) to the bulk external connection (millimeter scale). The flexible sheet electrode was a multilayered structure of the nanofilm, conductive lines, and insulating layers (Figure 1a,b and Figure S1). The substrate of the sheet electrode was a poly(D,L-lactic acid) (PDLLA) nanofilm (thickness: ∼400 nm; Young’s modulus: ∼4 GPa).11 Conductive lines were inkjetprinted onto the nanofilm, using poly(3,4-ethylenedioxythiophene):polystyrenesulfonate) (PEDOT:PSS) ink as an inkabsorbing layer, which were reported to have a physiological stability and little toxicity to neuronal tissues,12 and Au nanoink (total thickness of a conductive line: 122 nm, 42 nmthick Au lines and 80 nm-thick PEDOT:PSS lines). The Au nanoink was repelled by the pristine nanofilm (Figure S2a). On the other hand, the PEDOT:PSS ink was printed without ink pooling or breaking (Figure S2b); thus, the Au nanoink was inkjet-printed onto PEDOT:PSS lines (Figure S2c). The surface of the inkjet-printed conductive part became rough because of the Au nanoink consisted of aggregations of Au nanoparticles (tens to hundreds of nanometers in diameter) (Figure S2c).13 Then the printed nanofilm was dipped in a

electronic devices that are injectable and preassembled would be ideal. Polymeric thin films (referred to as “nanofilms”) are characterized by a thickness of tens to hundreds of nanometers and a large aspect ratio (over 106 between thickness and width).9 Owing to their features, free-standing nanofilms have notable flexibility and physical adhesiveness. Recently, we reported on flexible electronics made of nanofilms produced with drop-on-demand inkjet printing. Inkjet printing can draw various patterns with line widths as small as tens of micrometers using various inks such as conductive inks, so this technique has attracted attention as a way to form electronic circuits and useful for proof-of-concept before microfabrication methods.9 We also found that the inkjetprinted conductive patterns prepared on flexible nanofilms conformed well to skin, owing to the physical adhesiveness of the nanofilms.10 Therefore, using a nanofilm is a unique approach to integrate electrical elements on biological tissues in a flexible and conformable way. Here, we show a flexible neural electrode consisting of thin layers with an overall thickness of 6 μm that allows for in vivo recordings not only of spontaneous but also of optogenetic neuronal signals. This nanofilm-based electrode (referred to as a “sheet electrode”) was designed to solve the mechanical mismatch between the probes and the rigid external devices or the soft brain tissues, which fabricated as follows (see also Supporting Information). In brief, (i) 50-μm-wide conductive lines were inkjet-printed on the nanofilm, which was dipcoated with fluoropolymer to make insulating layers. (ii) The tip of the insulated nanofilm was trimmed with a razor, which exposed the microelectrodes. (iii) The resulting flexible film electrode was twisted from a 2D sheet into a 3D needle, which allowed it to be inserted into brain tissues. The electrode was B

DOI: 10.1021/acsabm.8b00574 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

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ACS Applied Bio Materials

the “needle electrode”) in a 5 mg mL−1 agarose gel, whose softness is similar to brain tissues.6,20 The needle electrode was inserted into the hydrogel, whereas the unprocessed one was not (Figure 2c). When the needle electrode was removed, 1 h after being inserted into the hydrogel, the electrode was unsolidified and became flexible again (Figure 2d). The solidified needle electrode turned soft again in the hydrogel because the PVA was dissolved, which would be expected in brain tissues. Twisting the sheet into a needle-like structure minimizes the tip size of the electrode, and the water-soluble polymer injected into the needle electrode increases its mechanical stiffness temporarily, which allows the electrode to be vertically inserted into the brain. The terminal edge connected to external devices had a different layered structure from the electrode side: it was composed of a printed nanofilm, a polyimide film (Young’s modulus: 3.3 GPa; thickness: 12.5 μm),21 and a precut tip of a flexible, flat cable (FFC; thickness: 200−300 μm) (see also Figure 1b). To connect the sheet electrode (bending stiffness: 22 nN m) to the rigid external device (over several hundreds of N m), we gradually increased the thickness and stiffness of the connection. Because polymer nanofilms had the physically adhesive property originated from the ultrathin structure, which was previously investigated using a micro scratch tester,9 the polyimide film was well attached to the printed nanofilm without glue and not separated from the PDLLA nanofilm (Figure S1a, (iii)). Thus, the bending stiffness increased from 22 nN m to 560 nN m, and the total thickness changed from ∼6 μm to ∼13 μm. The precut FFC was then attached to the polyimide film with double-sided tape (bending stiffness: more than several mN m), which increased the total thickness up to ∼313 μm. Finally, the preassembled connection, composed of the polyimide film and the precut FFC, was easily clipped with an FFC conversion board adapter in one step (Figure 2e). Gradual increment of the stiffness and thickness moderated the mechanical mismatch between the printed nanofilm and the rigid external device. The flexible and robust nanofilm had an array of conductive lines, and the adhesiveness facilitated the attachment of the nanofilm to a rigid substrate, which allowed us to increase the hardness gradually from the nanofilm to the external device. Notably, our process needs no extra process, such as vacuum and high-temperature or heating processing, to connect the flexible electronic probes to the external device after inserting the probes because the preassembled sheet electrode was already connected to the board adapter. The cut and exposed microelectrodes were observed by scanning electron microscopy with energy dispersive X-ray spectroscopy (SEM-EDS) (Figure 3a−d). The ∼50 μm-wide microelectrodes (total thickness of a conductive line: 122 nm) were clearly exposed on the cutting edge. The layered structure, CYTOP/Au/CYTOP, was maintained without crushing after cutting. Interestingly, while the conventional neural electrodes were round or square with an aspect ratio of ∼1, the exposed ultrathin microelectrodes showed a unique shape with an aspect ratio of >50 between thickness and width, which allowed for fabricating thin probes (∼6 μm). Then the electrical properties of the ultrathin microelectrodes were analyzed with electrochemical impedance spectroscopy (EIS) in the frequency range of 10 Hz to 100 kHz, which is the relevant frequency of neuronal activities.22 EIS produces a Bode plot of the impedance magnitude and phase angle, and it reveals the capacitive and resistive

solution of hydrophobic fluoropolymer CYTOP (1.5 GPa)14 to form an electrically insulating layer on both sides (thickness of one CYTOP layer: ∼2.8 μm) (Figure S1b). The resulting structure had a total thickness of ∼6 μm (Figure 1c; full details of the fabrication methods in Figure S1a). We calculated a bending stiffness, which is proportional to the Young’s modulus and the cube of thickness, because not only the Young’s modulus of materials, but also the thickness of the probe affects the flexibility of the probe. The bending stiffness of the PDLLA nanofilm and conductive line was 0.015 nN m11 and 0.0006 nN m,10,15 respectively,8 and the printed nanofilm (0.015 + 0.0006 = 0.0156 nN m) showed flexible movement without breaking (Supplementary Movie S1). The bending stiffness of the printed nanofilm coated with CYTOP (thickness, ∼6 μm; referred to as the “CYTOP-coated nanofilm”) was 22 nN m. This value was 7−2770-times smaller than that of typical neural probes such as wire tetrodes (6.1 × 104 nN m),16 silicon probes (4.6 × 104 nN m),17 carbon fiber (3.9 × 103 nN m),18 and polyimide-based probes (0.16−1.3 × 103 nN m)19 (according to Lieber group) (Tables S1 and S2). Such a small bending stiffness means that the CYTOP-coated nanofilm is much more flexible than conventional electrodes. In addition, the printed nanofilm is too flexible, making it difficult to handle in air and use in practice. Therefore, the CYTOP layers made the printed nanofilm not only insulated, but also robust. In fact, the CYTOP-coated nanofilm maintained its shape even when touched and twisted (see also Figure 1a). Next, simply by cutting with a razor, we exposed microelectrodes that could record neuronal signals (Figure 2a). We took advantage of the flexibility and robustness of the

Figure 2. Concept of the flexible polymer nanofilm electrode. (a) Micrograph of the electrode part cut out with a razor. (b) Micrograph of the electrode part of the needle electrode. The needle electrode was fabricated by twisting a sheet electrode, solidifying it with 10 mg mL−1 PVA, and cutting it with a razor. (c) Inserting the sheet electrode (left) and needle electrode (right) into 5 mg mL−1 agarose gel. The sheet electrode slipped on the surface and was not inserted. (d) Needle electrode, 1 h after insertion into the hydrogel, showed morphological transformation from rigid to flexible. (e) Photograph of the overall setup, with the black arrow indicating the mechanical support with a precut 200-μL pipet chip.

sheet electrode, twisting it three times to make a helical needle (Figure 2b; Figure S3a). Then we injected a 10 mg mL−1 poly(vinyl alcohol) (PVA) solution into the needle, dried it to solidify its shape, and cut a tip of the needle (Figure S3b). We verified the dissolution of PVA and the morphology of the needle-shaped sheet electrode after cutting (referred to as C

DOI: 10.1021/acsabm.8b00574 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

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ACS Applied Bio Materials

Figure 3. Structure and electrical properties of microelectrodes. (a) SEM cross-sectional image of one ultrathin microelectrode (aspect ratio of >50) and explanation of the detailed structure. (b) Magnified image of a dashed red box in panel a. (c,d) EDS element maps of panel b. (c, d) Fluorine and gold, respectively. (e, f) Bode plots summarizing the average impedance spectra in magnitude and phase of the sheet electrode and needle electrode, respectively. The error bars and the measures of center represent the standard deviation and the mean, respectively (n = 12 different electrodes).

properties of the microelectrodes. Bode plots summarized the average impedance and phase of sheet electrodes (12 microelectrodes individually) (Figure 3e) and those of needle electrodes (12 microelectrodes) (Figure 3f). The impedance magnitude of the sheet electrodes was 160−340 kΩ at 1 kHz. After twisting the sheet electrodes and cutting the tip of the twisted needle electrodes, the impedance magnitude of the needle electrodes was 150−600 kΩ. The standard deviation of impedance became larger after twisting, while those impedance values were in the same range as those used for extracellular recording.23 In the case of sheet electrodes, all the microelectrodes were exposed by vertically cutting with a razor with respect to the wiring direction, and there is little stress of the conductive parts. For the needle electrodes, by contrast, the razor seemed to tilt the edge of the needle electrode because the conductive wirings were twisted, so the exposed position and area of each microelectrode would be slightly changed, which resulted in the increment of the standard deviation in the recorded impedance. In addition, the conductive part received some strain after twisting and then the resistance increased. Although the cutting after twisting was the stressful operation, the electrode showed the impedance range of 100

kΩ to 7 MΩ without breaking when the cutting and measurement were repeated three times (data not shown). Then we also investigated the phase angle. The phase angle of the exposed sheet electrodes printed with PEDOT:PSS and Au nanoink was −40° to −10°, which indicates resistive electrodes at physiologically relevant frequency of 1 kHz.22 On the other hand, bare PEDOT:PSS sheet electrodes without Au nanoink showed a phase angle of −30° to −10° (Figure S2d). This behavior, the similar range of phase angle for PEDOT:PSS and Au nanoink electrodes, inspires that the resistive charge transfer at an electrolyte−electrode interface was not affected after printing Au nanoink over the PEDOT:PSS layer. After fabricating the sheet electrodes, we assessed their ability to record neuronal activity. First, ex vivo recording experiments were conducted using a needle electrode on a transgenic rat brain slice expressing channelrhodopsin-2, which enabled neuronal excitation by illumination with blue light.24 Previously, we reported that optogenetic stimulation induced the striatal firing.25 The striatal slices were prepared by the reported methods.25,26 The tip of the needle electrode was attached to the surface of the prepared striatal slice. To photostimulate the striatum, we used epi-illumination from an D

DOI: 10.1021/acsabm.8b00574 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

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ACS Applied Bio Materials

Figure 4. Acute in vivo rodent brain recording. (a) Recording setup. The needle electrode was inserted into a mouse thalamus. (b) Representative spontaneous activities recorded in the thalamus at a depth of 2.2 mm from the cortical surface. The right side shows a magnified waveform. (c) Photograph of the overall setup. The unwrapped flexible part was connected to the external device to the side. (d) Micrograph of the electrode part of the nanofilm optrode. The optrode was fabricated by wrapping a sheet electrode around an optical fiber (diameter of 500 μm), fixing it with 10 mg mL−1 PVA, and cutting it with a razor. (e) Setup for optogenetic stimulation and recording of evoked in vivo neuronal signals using a channelrhodopsin-2 expressing Wistar rat. The optrode was inserted into a rat hippocampus. (f) Representative optogenetically evoked signals at a depth of 3.3 mm (the magnified spike in the right panel is indicated by a red arrow in the left panel). Photostimulation (blue bar) with a duration of 1 s evoked spikes with various amplitudes.

reference electrode was placed on the dura mater. The needle electrode was inserted into the thalamus using a micromanipulator (Figure 4a). One of four channels successfully recorded the spike signals (Figure 4b). The observed maximum spike had a negative peak at −62.7 μV. The standard deviation (SD) of the recorded baseline noise was 7.3 μV. Because the threshold criterion of appropriate spike detection is −4SD,26 the detected spike amplitude is sufficiently large. On the other hand, there are some reasons that electrodes did not work well. One or two of four electrodes had already broken due to breaking of conductive lines or had not worked well due to electrical disconnections with the external device in the fabrication process. In addition, the comparatively large diameter (200−300 μm) of the needle

upright microscope with a 470 nm LED. Because most of the striatal neurons are deeply hyperpolarized, spontaneous firings are not normally observed. By using the needle electrode, we successfully recorded the optogenetically evoked firings (Figure S4a), similar to the striatal firing pattern recorded with a nickel−chrome wire tetrode (Figure S4b). The optogenetically evoked striatal firing pattern of the transgenic rat showed characteristic properties: the number of firings during and after photostimulation increased progressively with repeated photostimulation, similar to the results of our previous report.16 Next, acute in vivo recording tests were performed on adult C57BL6 mice. The mice were deeply anesthetized with isoflurane and then stereotaxically head-fixed. A silver wire E

DOI: 10.1021/acsabm.8b00574 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

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spike signal sensing, and segmenting the electrode tips to precisely record in vivo neuronal activity.

electrode most probably reduced the chances of effective contact with neurons in the recording experiment. As a result, one of four channels recorded spikes. Finally, we demonstrated the utility of the sheet electrode for optogenetics on the rat expressing channelrhodopsin-2. The sheet electrode was integrated with an optical fiber (diameter of 500 μm) by wrapping the electrode around the fiber (Figure 4c,d; Figure S5). The electrode protruded beyond the fiber for ∼500 μm. Slack in the unwrapped portion of the optrode allowed us to connect it to the preamp connector and prevented the optrode from breaking during insertion (see also Figure 4c). The channelrhodopsin-2 expressing rat was deeply anesthetized with isoflurane and the stereotaxically head-fixed. The optrode was inserted in the rat hippocampus. Localized blue light evoked hippocampal action potentials with various amplitudes, which were limited to the photostimulation period (Figure 4e,f), similar to the results of our previous report.27 In conclusion, we developed an inkjet-printed, flexible neural electrode in which the ultrathin Au electrode was easily exposed by cutting off the tip, and solid spikes were detected by using the electrode. Taking advantage of the flexibility, robustness, and physical adhesiveness of the nanofilm, we created preassembled connections to rigid external devices. We also shaped the CYTOP-coated nanofilm into a needle allowing for insertion into the brain tissues and integration of the optical fiber for optogenetics. The limitation of the sheet electrode complex is that there are a few spike-detecting efficient electrodes out of the total. In general, an electrode receives multiunit spikes from several neurons surrounding the exposed electrode surface (for example, Figure 4b). To sort spike units from the multiunit raw data, efficient multichannel recording and spike sorting method are required.28 We estimate that the low efficiency of our needle electrode complex is due to the manual fabrication methods and large size of the probe tip (200−300 μm). It is our current goal to miniaturize the probe size by varying the electrode complex geometry to enhance the contact with neurons. We are currently trying to miniaturize the electrode tip by vertically staking the assembling process of the ultrathin electrodes (aspect ratio >50) and keeping total thickness thin to minimize invasiveness and increase the accuracy and reliability of the probe. Then we will observe/evaluate tissue damage by insertion of the electrodes and an interaction between biological tissues and the sheet electrodes to understand how the electrodes agglutinate with neurons and affect glial cells after insertion of the electrodes. Our probe design and fabrication technique are promising approaches for bridging ultrathin electrodes and bulky external connections. The present study is also the first proof-of-concept study to apply the ultrathin microelectrode fabricated on the flexible nanofilm in neural technology; the process will be useful not only for insertable neural probes but also for tissue-contact ECoG electrodes. Polymer nanofilms and inkjet-printing process allow researchers to develop insertable neural probes with various inks and patterning before fabricating sophisticated products with microfabrication processes. For example, the inkjet-printing process enables us to choose different inks such as metals, conductive polymers and carbon-based nanomaterials, which would be useful for further functionalization of the probes for electrochemical analysis.29,30 The flexible and robust structure of our probes will allow for tailor-made customization for various research requirements such as changing the complex geometry, increasing the number of electrodes for

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsabm.8b00574. Materials and methods, detailed materials characterizations (PDF) Supplementary movies (AVI)

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected]; [email protected]. Phone: +81-45-924-5712. Fax: +81-45-924-57. *E-mail: [email protected]. Phone: +81-42-995-1483. ORCID

Shinji Takeoka: 0000-0002-6230-1517 Toshinori Fujie: 0000-0003-1417-8670 Author Contributions

N.K. designed the neural probes, performed the experiments, and wrote the manuscript with the supervision of K.Y., S.T., and T.F. H.O. edited the manuscript and analyzed the results. M.A. helped with the recording experiments with the supervision of Y.M. and H.O. H.O. and T.F. supervised the whole project. Funding

This work was supported by the Precursory Research for Embryonic Science and Technology (PRESTO) program from the Japan Science and Technology Agency (JST; Grant No. JPMJPR152A), JSPS KAKENHI (Grant Nos. 16K08923, 17K20116, 18H03539, 18H05469), a Grant-in-Aid for JSPS Fellows (Grant No. 16J07140), JSPS Core-to-Core Program, the Noguchi Institute, and the Tanaka Memorial Foundation. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS The authors thank Assoc. Prof. Yosuke Okamura (Department of Applied Chemistry, Tokai University) for providing the information on insulating layer.

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REFERENCES

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DOI: 10.1021/acsabm.8b00574 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX

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DOI: 10.1021/acsabm.8b00574 ACS Appl. Bio Mater. XXXX, XXX, XXX−XXX