Treatment trends for Parkinson’s disease are revolutionizing from conventional pharmaceutical treatment depending on the patient’s symptoms to closed-loop electroceutical treatment in order to minimize the crucial pharmacokinetic drawbacks. Over the decades, closed-loop neuromodulations have been achieved by electroceutical devices based on continuous electrical stimulation and sensing of electrophysiological biomarkers. However, critical drawbacks have been observed with vulnerable to artefacts from the nearby stimulation electrode and frequent adjustment of stimulation parameters. On the contrary, direct monitoring of neurochemicals allows high spatial resolution, signal specificity and interference-free detection compared to closed-loop DBS. Here, we propose an artificial homeostasis system with electrochemical sensing and pharmaceutical stimulation for impaired brain function. In order to implement the artificial homeostasis system by mimicking the brain system, we have reported an implantable multi-deformable double-sided DA-sensing probe with an all-in-one structure by integrating three-electrode system and drug refillable microfluidic probe with controllable injection for synchronized diagnosis and treatment. The experimental and computational analysis clearly indicate that the proposed probes could serve as potential tools for building artificial homeostasis by the bio-responsive closed-loop system with an electropharmaceutical approach.
Article
Artificial homeostasis system with electrochemical sensing and pharmaceutical stimulation for impaired brain function
https://doi.org/10.21203/rs.3.rs-2417950/v1
This work is licensed under a CC BY 4.0 License
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electrochemical biosensor
drug delivery
soft mechanics
neural probe
biomedical device
chemical closed-loop
Parkinson’s disease
As the elderly population increases, the prevalence of neurodegenerative disorders is also increasing1–3. Among the neurodegenerative diseases, Parkinson’s disease (PD) is one of the most frequent neurological disorders worldwide4. The inadequate control of dopamine (DA) levels in the central nervous system (CNS) is associated with the pathogenesis of this prevalent neurological condition, as DA is one of the most important catecholamine neurotransmitters widely present in the CNS5. DA regulates several aspects of the brain circuitry, playing an important role in motivation, motor control, memory, reward, etc6–9. Therefore, as a powerful neurotransmitter, analysing and controlling DA levels in the brain is critical for neurodegenerative disorders, particularly PD.
Over the last half-century, a wide range of therapies have been accessible for PD patients, with the majority relying on drugs and physical therapy to control symptoms10. Among various treatments for PD patients, oral administration of levodopa (L-DOPA) still remains the gold standard due to its impressive efficacy on motor symptoms and its relatively affordable price compared to other available medications11,12. Despite these key benefits of oral administration of L-DOPA, critical pharmacokinetic drawbacks such as excessive oscillations in the DA level, motor fluctuation and dyskinesia that are caused by intermittent stimulation of striatal DA receptors and dysfunctional dopaminergic pathways, make long-term treatment difficult13,14. To minimize the pharmacokinetic side effects caused by oral administration of L-DOPA, many studies have been conducted on intracerebroventricular (ICV) administration of DA15,16. This approach has risen to present the evidence as a working concept for entire stages of PD by reducing motor impairment and boosting DA levels in the brains of animals17,18 and PD patients19,20. Despite the various treatment methods, such as oral administration of L-DOPA and ICV administration of DA, limitations and difficulties in treating PD patients still remain as these treatments are performed as open-loop neuromodulation, in which qualitative behavioural indicators trigger drug administration. Additionally, recent studies have shown closed-loop neuromodulation for neurodegenerative disease treatment allows substantial clinical benefits and effectiveness compared to open-loop neuromodulation21,22.
Closed-loop neuromodulation has traditionally been achieved through deep brain stimulation (DBS) based electroceutical devices, which employ action potentials (AP) and local field potential (LFP) as electrophysiological biomarkers23,24. Closed-loop DBS is a therapeutic strategy based on continuous electrical stimulation with high frequency (typically 100–180 Hz) using chronically various implantable electrodes25–27 that has proven to be an effective therapy for PD patients through electrical stimulation of the subthalamic nucleus (STN)28–30. Although the closed-loop DBS has shown therapeutic effects for symptomatic relief in PD, this approach is highly vulnerable to artefacts from the nearby stimulation electrode since it detects electrical signals from surrounding neurons31,32. Owing to these artefacts, closed-loop DBS devices with neurochemical analysis techniques, such as amperometry33,34 and voltammetry35,36, have emerged to directly detect local changes in neurotransmitter concentration that are evoked by internal and external electrical stimuli as feedback signals. Neurochemical monitoring allows real-time characterization of neural activity with high spatial resolution, signal specificity and interference-free detection from other electrophysiological signals compared with closed-loop DBS with electrical neural recordings37,38. Despite current efforts to construct closed-loop systems to examine and treat neurodegenerative disease in real-time, the fundamental operating mechanism induced by DBS is not fully understood due to limited information39,40. Besides, the trial-and-error process is currently used to sustain therapeutic benefits with stimulation-induced side effects by subsequently adjusting several parameters such as electrode contacts, stimulus frequency, amplitude, and pulse duration every few months41,42. Electropharmaceutical closed-loop system synchronized with biological system for homeostatic control, on the other hand, can be employed for chronic treatments of neurodegenerative diseases by utilizing neurotransmitter concentration levels as biomarkers to modulate brain function with neurochemical drug infusion43–45. However, challenges still remain unsolved to provide the electropharmaceutical closed-loop system with controllable and precise drug injection along with continuous monitoring of DA. Currently, many clinical-grade neurological probes and drug injection devices are fabricated in rigid form, which is incapable of chronical sensing and stimulation due to its mechanical incompatible with soft and curvilinear brain tissue46–48. In an environment such as the brain, a flexible probe is essential, since mechanical mismatch with soft neural tissues induced by a rigid probe can hinder long-term functionality and restrict anatomical applicability. Therefore, flexible electronics for an electropharmaceutical closed-loop system are required, but such probes have rarely been reported.
Here, we introduce a novel chronic implantable multi-deformable double-sided (MDD) neural probe and ICV administrative microfluidic probe for a synchronized DA-sensing and stimulation system to implement artificial homeostasis for neurodegenerative disorders, due to the importance of DA in the brain circuitry. To utilize the MDD DA-sensing probe and microfluidic probe as the synchronized DA-sensing and stimulation system for electropharmaceutical treatment, new materials design and heterogeneous integration of multiple elements are introduced. Firstly, the proposed DA-sensing probe is developed with highly stretchable materials for intimate integration with the brain tissue and chronic monitoring with minimizing undesirable noise which may cause by motion artefacts. Furthermore, for minimal damage to the brain tissue during real-time monitoring of DA, the MDD probe was heterogeneously integrated with three different electrodes, such as a working electrode (WE), reference electrode (RE), and counter electrode (CE), in a single unit probe through a double-sided fabrication process. As opposed to the conventional electrochemical methods in which two or more holes were drilled into the skull to apply electrochemical measurement49–51, the MDD DA-sensing probe requires a single drill that provides relatively less invasive implantation. In addition, high DA sensitivity was achieved through enzyme immobilization on the 3-dimensional nano-forest-like structures grown on the surface of WE, which maximized the surface area that effectively enhanced the DA sensitivity. Secondly, for long-term chemical stimulation into the deep brain, the injectable microfluidic module consists of: (1) an electrically driven micropump structure with thermos-pneumatic actuation for the administration of an adequate amount of drug dose, (2) a drug reservoir that can replenish the drug repeatedly, and (3) programmable two-dimensional microheater array for drug administration with precise control of target flow rate, which is necessary to prevent deviation of DA level from the normal range due to the excessive pharmaceutical treatment. In this study, we demonstrated the capability of controllable operation of a microfluidic probe and real-time measurement of the DA-sensing probe for in vivo. These proof-of-principle studies of continuous and deterministic methods will open the route to a minimally invasive artificial homeostasis system based on a closed-loop platform as neurochemical sensing and stimulation engineered as a quantitative index for neurodegenerative disorders.
Artificial and homeostasis system as electropharmaceutical treatment with DA-sensing probe and microfluidic probe for impaired brain function. PD is a highly prevalent neurodegenerative disease caused by the dysregulation of DA in the brain. Despite the introduction of a variety of treatments for PD, most of the treatments are symptomatic. Traditionally, oral administration of L-DOPA remains the most potent pharmaceutical treatment for controlling PD symptoms even though it is associated with motor complications, such as motor fluctuations and dyskinesias. In order to minimize pharmacokinetic issues observed during the treatment with oral administration of L-DOPA, the soft and miniaturized DA-sensing probe and the microfluidic probe utilizing directly into the brain as ICV administration is a potential new treatment of neurodegenerative disease by mimicking the physiological released of DA and actively controlling the DA levels in the brain. As shown in Fig. 1, utilization of an artificial homeostasis system probes for PD, DA-sensing probe and microfluidic probe, are implanted at caudate putamen (CPu) in the left hemisphere of the brain with degenerated dopaminergic axons in which anatomically located at the nigrostriatal dopaminergic pathway to monitor the levels of DA (bottom right inset) and lateral ventricle (VL) on right hemisphere as an ICV route for DA infusion in the brain (bottom left inset). The DA-sensing probe which consists of three electrodes, such as WE, RE, and RE, monitors the signals of the DA level in real-time through analogue signal processing units that include a filter, buffer, amplifier, and analogue-to-digital converter (ADC). The signals representing the DA level in the brain induce ICV administration of DA into VL as programmed by manipulating the operation of the micropump through MC and other integrated circuits (ICs), such as general-purpose input/output (GPIO) and switch ICs (upper left inset).
A multi-deformable double-sided (MDD) DA-sensing probe design. Figure 2a is an exploded schematic view of the MDD DA-sensing probe to illustrate the layer information and associated components of the DA-sensing probe. The MDD DA-sensing probe consists of three primary components: the top component (Fig. 2a- I), the bottom component (Fig. 2a-III), and a stretchable substrate (Fig. 2a-II) of poly(dimethylsiloxane) (PDMS; ~ 250 µm in thickness) to combine micrometer-scale three electrodes (WE, RE, and CE) in one single probe to perform a reliable electrochemical analysis. The fabrication processes for top and bottom components to implement a three-electrode system begin with spin casting and consecutive thermal curing of poly(methylmethacrylate) (PMMA; ~ 1 µm in thickness) as a temporary sacrificial layer and photocurable epoxy polymer (SU-8; ~ 5 µm in thickness) as a dielectric layer on a glass substrate, followed by photolithography patterning. Next, bi-layer of Cr/Au (7 nm/200 nm in thickness) and Ti/Pt (20 nm/100 nm in thickness) are deposited by electron beam evaporation on the separate samples undergoes photolithographic patterning to create WE/RE (~ 0.07 mm2/~0.07 mm2) for top component and CE (~ 0.55 mm2) for bottom component, respectively. The ZnO is deposited by RF-sputtering deposition and photolithographic patterning creates a seed layer only on WE surface to develop ZnO nanorods (NRs). ZnO NRs are grown on micrometer surface of WE to function as 3D matrixes for enzyme immobilization by crosslinking (CL) process, which allows high sensitivity by maximizing the surface area of the WE. An additional layer of dielectric polymer is formed via spin casting and curing, followed by photolithography patterning. The fabrication process is completed by integrating top (WE/RE) and bottom (CE) components on each side of the prepared PDMS substrate in virtue of strong siloxane bonding via oxygen plasma treatment. For the DA detection, the WE have selectively coated with a highly sensitive and selective enzyme (Tyrosinase, TYR; EC 1.14.18.1 2840 U mg− 1) layer, and the RE is coated with silver/silver chloride (Ag/AgCl) paste. Further fabrication details of the MDD DA-sensing probe can be found in the experimental section and Supplementary Fig. 1,2. Figure 2b shows an overall representative image of the fabricated MDD neural probe (upper left inset), microscopic image to MDD neural probe tip with WE/RE on one side of the probe and the other side with CE (upper right inset), and optical image of MDD neural probe implanted on a phantom brain (0.6% agarose gel), showing minimal invasive insertion (bottom).
Figure 2c shows schematics of the MDD DA-sensing probe’s WE surface with immobilized enzyme (TYR) on hydrothermally grown ZnO nanorods (NRs) with the reaction scheme for DA detection. The detection mechanism of DA is based on the oxidation of DA into o-dopaquinone (DQ) (Eq. (1)), and the subsequent electrochemical reduction of DQ on the surface of WE at low potential (Eq. (2))52.
o-Dopaquinone + 2H+ + 2e− → Dopamine + H2O (2)
The immobilization of enzyme (TYR) is crucial step in the fabrication of the MDD DA-sensing probe for the enzymatically catalyzed oxidation of DA. For the effective immobilization of TYR on WE with larger surface area to increase the reduction current for higher sensitivity, the 3D nano-forest-like structures are utilized by hydrothermally grown ZnO NRs on the WE surface of the MDD DA-sensing probe. The development of ZnO NRs is well explained in the experimental section.
To immobilize the enzyme on the WE of the probe, the hydrothermally grown ZnO NRs (Supplementary Fig. 3a) are coated with silica to form ZnO-O-Si, which hydrolysed tetraethyl orthosilicate (TEOS, precursor of silica) condensed on the surface of ZnO NRs53. After the silanisation process, 3-(Aminopropyl)triethoxysilane (APTES, amino-modification reagent) is added to the reaction solution to form amine-functionalized surface of ZnO NRs (ZnO@SiO2-NH2) by sol-gel process (Supplementary Fig. 3b). After amine-functionalization of ZnO NRs, the TYR is immobilized on the amine-modified ZnO NRs structures by CL with glutaraldehyde (GA), which acts as a bridge with aldehyde group between primary amino groups (-NH2) of silanized ZnO NRS and the terminal amino groups of the TYR (Supplementary Fig. 3c)54. The heteroarchitecture of surface modified WE is confirmed from scanning electron microscopy (SEM) and energy dispersive X-ray (EDX) analysis, as shown in Fig. 2d-f and Supplementary Fig. 3. From the SEM images with enlarged magnification, densely populated ZnO NRs having the shape of a hexagonal pillar is clearly observed. The EDX spectrum confirms the presence of well-grown ZnO NRs showing strong intensity oxygen (O) and zinc (Zn) peaks (Fig. 2d). In addition, the surface modification processes for amine-functionalized WE with TEOS and APTES are verified. As shown in the SEM image (Fig. 2e), ZnO@SiO2-NH2 structures are formed, and it is verified that the surface of the ZnO NRs became thicker. Also, unlike the EDX analysis data of ZnO NRs, the additional intensity peaks of silicon (Si), nitrogen (N) and carbon (C) are observed, indicating structural changes of ZnO NRs due to the interaction with TEOS and APTES, which corresponds to the silanization process and amine-functionalization process, respectively. The enzyme immobilization process is also investigated by SEM image and EDX analysis (Fig. 2f). From the SEM image, the amine-functionalized ZnO NRs have linked each other with the enzyme (TYR) in virtue of CL process with GA, and compared to previous processes (Fig. 2e), the stronger intensity of C and O peaks are examined (Supplementary Table 1), indicating successful immobilization of TYR onto the amine-functionalized ZnO NRs.
The electrochemical properties for each surface modification process of the fabricated MDD DA-sensing probe are investigated by cyclic voltammetry (CV) method in 0.05M phosphate buffer saline (PBS) solution (pH 7.4) between potentials of -0.5 V and 0.5 V at a scan rate of 150 mV/s. CV was utilized to compare electrochemical reactions on the surface-modified electrodes, and CV measurements of the traditional planar Au, Au/ZnO seed layer, Au/ZnO NRs, and Au/ZnO NRs/CL/TYR electrodes were taken to study the successive modification process of DA-sensitive MDD neural probe, as shown in Supplementary Fig. 4. The CV curve of the bare Au electrode (curve I) shown the largest current response compared to the CV curves of other modified electrodes, such as Au/ZnO seed layer (curve II) (ZnO; ~ 7.261 x 107/ cm∙Ω in conductivity)55, Au/ZnO NRs (curve III), and Au/ZnO NRs/CL/TYR (curve IV), due to the excellent electrical conductivity (Au, ~ 0.452 x 106/cm∙Ω). The hysteresis loop areas of CV start to decrease after the bare Au electrode modification process advance, due to the increased electron transfer resistance, causing the potential drop across the surface coating and the resulting reduction of electrical conductivity. Furthermore, the electrochemical characteristic of CV confirms the larger surface area of WE owing to the ZnO NRs (curve III) by presenting a reduction in current response compared to the electrode with ZnO seed layer (curve II). The redox current of the enzyme immobilized electrode (TYR, curve IV) on ZnO NRs significantly decreased due to the limited conductivity of the enzyme (TYR), indicating the successful immobilization of enzyme onto the WE of MDD DA-sensing probe.
Moreover, the electrochemical properties of the fabricated MDD DA-sensing probe with (blue curve) and without (red curve) ZnO NRs are compared in the absence (dotted line) and presence (solid line) of 100 µM of DA (Fig. 2g). As shown in the Fig. 2g, the redox peak of the electrode with ZnO NRs in the absence of DA (blue dotted curve) barely shown any obvious current peak. However, when the CV is measured in the presence of 100 µM DA, the electrode with ZnO NRs (blue solid line) shows the largest current response compared to that of the electrode without ZnO NRs (red solid line), and a notable reduction peak (peak potential of -330 mV) is observed due to the electrocatalytic reduction of DA. The enhanced current response can be attributed to the high aspect ratio of ZnO NRs, which provide a high specific area for enough enzyme loading on hydrothermally grown ZnO NRs. The amperometric response of the fabricated MDD DA-sensing probe (Au/ZnO NRs/CL/TYR) are conducted to determine the performance of the DA-sensing application by successively adding DA in an increasing concentration range of 1 to 9 µM, in 0.05M PBS solution (pH 7.4) at an applied potential of -330 mV (vs. Ag/AgCl) under stirring condition as shown in Fig. 2h. The MDD DA-sensing probe showed excellent amperometric response with successive addition of 1 µM DA in 0.05M PBS solution (pH 7.4). A well-defined amperometric response presented a short response time of ~ 1s.
The anti-interference properties are critical parameters for electrochemical sensing, as the working potential may be related to the reduction of other potential interfering species present with DA among other biochemical analytes, including ascorbic acid (AA) and norepinephrine (NE), whose reduction potential and molecular structures are similar to that of DA, respectively. Also, better selectivity will ensure high accuracy during measurement. As shown in Fig. 3i, the temporal measurements of amperometric responses were assessed by successive addition of 500 µM DA, and 100 µM interfering species i.e., AA and NE, respectively. Adding DA resulted in a significant and rapid current response, whereas adding interfering species (AA and NE) caused negligible current variations. Compared current response for DA detection with other interfering species was illustrated by the histogram in Fig. 2j. When the amperometric response of 100 µM DA was set as 100%, the graph shows that the interfering species caused a very low response i.e., for 100 µM L-DOPA (~ 35%), for 500 µM NE (~ 2.5%), and for 500 µM AA (~ 0.7%). The inset of Fig. 3e presents the current response for interfering species addition of AA and NE. Additionally, the long-term stability of the MDD DA-sensing probe was also evaluated, and the related histogram in Fig. 2k showed good stability over 4 weeks, which retained approximately 90% of the initial sensitivity, revealing its outstanding stability. To assess the brain biocompatibility of the MDD DA-sensing probe, representative confocal fluorescence images from cultured catecholaminergic cells (SH-SY5Y cells)56 on the MDD DA-sensing probe showed a statistically negligible difference in cell viability when compared with control up to 7 days (Fig. 2l, Supplementary Fig. 5a-d). Also, as shown in Supplementary Fig. 5e, the MDD DA-sensing probe observed excellent biocompatibility on catecholaminergic cells with a negligible statically difference compared to the control sample.
Computational stress analysis of the MDD DA-sensing probe in complex deformation. The low elastic modulus of the materials used for brain implantable probes plays an important role in ensuring the reliable functionality of the probe. In a recent study, a soft polymer such as PDMS or polyimide (PI) gets much attention as promising material for the implantable probe to minimize the tissue damages58,59. The mechanical mismatch between thin metal film, which is normally employed for electrodes, and soft polymer substrates, however, results in mechanical instability, for the thin metal layers of electrodes have a relatively small elastic limit (strain of 1 ~ 4%) unlike the soft materials. Thus, the electrodes on both side of the probe are more vulnerable to this issue, since the excessive normal stress in the x-direction induced by the bending is proportional to the distance between the neutral plane and the electrodes according to the following Eq. (3),
σxx = κEzd (3)
where σxx is the normal stress in the x-direction, κ is the radius of curvature, E is elastic modulus, and zd is the distance between the electrode and the neutral plane. Therefore, the double-sided structure tends to be ruptured under even small longitudinal strain, as the electrodes are placed far from the neutral plane, which is closely located within the centre of the flexible substrate of the probe. In order to overcome this intrinsic limitation of the double-sided structure for the probe, the electrodes interconnections are designed in serpentine shape rather than conventional straight-line to allow longitudinal strain up to large scale under various deformation, such as simple deformation (i.e., stretching, bending) and complex deformation (i.e., buckling, twisting, and rolling).
The stability of the probe is verified by the finite element analysis (FEA) comparing conventional electrodes with straight-line (STL) and electrodes with serpentine-line (SPL) (Fig. 3a-f). The simulation parameters and models of the FEA are detailed in Supplementary Table 2 and Supplementary Note 1, respectively. When the probe is stretched in the x-direction for 30 µm, the average stress in the STL electrodes is 302 MPa and 785 MPa for the top electrodes (WE/RE) and bottom electrode (CE), respectively. On the other hand, the average stress in the SPL electrodes is 25.9 MPa and 39.7 MPa for WE/RE and CE, respectively, which is extremely low compared to that in STL electrodes (Fig. 3a). The graph represents the normalized stress as ultimate strength of three electrodes (WE, RE, and CE) comparing between STL and SPL to highlight the material failure (yellow regime) of the thin metal layers. The material failure is easily occurred at STL electrodes, with elongation at a few micrometers. In contrast, SPL electrodes are mechanically stable when stretched over 50 µm. In addition, when the probe is bent in the y- and z-directions, the stress exhibits the same behaviour as when the probe is stretched (Fig. 3b,c). When the y-direction force applied to the probe tip causes a 1 mm displacement from its initial position, the average stress in SPL electrodes (WE/RE ~ 29 MPa, CE ~ 48 MPa) is notably smaller than that in STL electrodes (WE/RE ~ 171 MPa, CE ~ 469 MPa), which resulting material failure on STL electrodes as opposed to SPL electrodes (Fig. 3b). The average stress over displacement (2 mm) on z-direction is observed as shown in Fig. 3c. According to the FEA simulation, the average stress in the SPL electrodes (WE/RE ~ 11 MPa, CE ~ 60 MPa) is ~ 9 folds less than the average stress in the STL electrodes (WE/RE ~ 99 MPa, CE ~ 253 MPa) despite the electrodes are far from the neutral plane.
Furthermore, as shown in Fig. 3d-f, the average stress in the STL and SPL electrodes during the complex deformation, such as buckling (Fig. 3d), twisting (Fig. 3e), and rolling (Fig. 3f), is analysed. Similar tendencies are observed to that of simple deformations. Figure 3d depict the buckling stress of the probe that generally occurs while inserting a probe into the brain. When the probe is compressed by 2.2 mm, the average stress showed significant differences between STL electrodes (WE/RE ~ 336 MPa, CE ~ 521 MPa) and SPL electrodes (WE/RE ~ 36 MPa, CE ~ 86 MPa). The SPL electrodes remain intact with 3 mm compression; however, the STL electrodes are easily damaged when the probe is buckled (compression of ~ 100 µm). According to the previous analysis, the SPL electrodes ensure mechanical stability without causing probe damage when buckling stress is given to the electrode during insertion into the brain. To highlight the outstanding flexibility of the MDD DA-sensing probe, the twisting deformation is analysed with rotation of 360° as shown in Fig. 3e. The average stress in the SPL electrodes (WE/RE ~ 112 MPa, CE ~ 166 MPa) is much smaller than that in the STL electrodes (WE/RE ~ 207 MPa, CE ~ 649 MPa). Also, the average stress induced by the rolling deformation is evaluated via twisting the probe 360°, followed by compressing the probe along the x-direction (Fig. 3f). When the probe is compressed by 1.8 mm during 360° twisting, the average stress in the SPL electrodes (WE/RE ~ 54 MPa, CE ~ 113 MPa) shows a significantly lower value than that in STL electrodes (WE/RE ~ 432 MPa, CE ~ 695 MPa). All the aforementioned stress analysis results demonstrate that the SPL electrodes are mechanically stable compared to the STL electrodes. Despite the mechanical mismatch between the thin metal film of the electrodes and soft polymer of the substrate, which are the main components of the MDD DA-sensing probe, the proposed design of heterogeneous integration with SPL electrodes ensures the reliability of the all-in-one DA-sensing probe even during the undesired deformation.
In Fig. 3g-I and Supplementary Fig. 6, SEM images of MDD DA-sensing probe are presented with various deformation modes: buckling (Fig. 3g and Supplementary Fig. 6a), twisting (Fig. 3h and Supplementary Fig. 6b,d), and rolling (Fig. 3i and Supplementary Fig. 6c), highlighting robust integration of the three essential electrodes (WE, RE, and CE) on a stretchable PDMS as an all-in-one electrochemical neural probe without any damage induced on the electrodes in virtue of the double-sided structure of the probe under dynamically deformed conditions, which minimizes the brain tissue damages normally associated with the conventional rigid neural probe.
Design and operation of microfluidic probe with peristaltic pumping. Figure 4a,b present the schematic diagram and optical images of the microfluidic probe with peristaltic micropump, respectively, which highlight the pump architecture to take the external air inside and push internalized drug out to the targeted brain region (detailed fabrication process available in the Supplementary Fig. 7). This drug delivery system is composed of four functional layers: (1) drug reservoir, which stores and refills the drug up to tens of µL, (2) microfluidic layer with channel (100 x 100 µm2) for drug delivery, (3) pumping layer for peristaltic operation via sequentially thermal expansion of three actuation chambers and (4) glass substrate patterned with planar electroresistive microheaters (Cr/Au; 7 nm/200 nm in thickness). The peristaltic operation in the air pathway above the micropump induces the delivered air into the reservoir to push out the drug. Functional parts such as microheaters, peristaltic micropump and reservoir are located above the skull and the injectable probe part (~ 500 µm width, 10 mm length) of the microfluidic layer can be inserted in the targeted brain region (Fig. 4b). As shown in SEM images (Fig. 4c), the pumping layer consists of three actuation chambers with deformable diaphragm membrane (Fig. 4c, top), which are deterministically aligned with microheaters to activate each diaphragm membrane independently (Fig. 4c, bottom right). When an actuation chamber is inflated by Joule heating, the elevation of the diaphragm membrane induces positive air pressure inside the microfluidic layer (Fig. 4d, top and Supplementary Fig. 8). Conversely, recovery of the diaphragm membrane due to cooling induces negative air pressure inside the microfluidic layer, which retakes the external air in to balance out the pressure difference (Fig. 4d, bottom). Figure 4e illustrates the concept and thermal images of device operation with peristaltic pumping using sequential thermo-pneumatic actuation (detailed thermal analysis available in Supplementary Fig. 9). Initially, all actuation chambers set inflation state to reach maximum positive air pressure in the microfluidic layer (Fig. 4e, step 1). Then a recovery of the diaphragm membrane connected to the air inlet provides the inflow of air from outside (Fig. 4e, step 2). The Sequential operations of the diaphragms control airflow direction in the micropump. Especially the re-expansion of the diaphragm recovered in the previous step makes the airflow direction to be controlled towards the reservoir (Fig. 4e, step 3,4). Cyclic operation of these four steps enables continuous and chronic drug delivery until the drug is fully consumed from the reservoir.
The effect of peristaltic pumping is affected by the design of microfluidic layer, whereby we determine design parameters that influences the flow. In terms of hydrodynamics, the variations in the radius of the pumping diaphragm membrane (R), the width of the microfluidic channel (W) and the angle of nozzle/diffuser structure of air pathway (θ) can affect the flow rate (Fig. 4f and Supplementary Fig. 10a). When assuming the amplitude of the diaphragm deflection does not vary along the radius of the diaphragm membrane, the volume change of the actuation chamber during deflection is proportional to the square of the radius of the diaphragm membrane57. With the increases of the diaphragm membrane radius (Supplementary Fig. 10b), net airflow rate also increases because of the increased volume change in air pathway, which raises the drug flow rate (Fig. 4g). Additionally, the width of the microfluidic channel (Supplementary Fig. 10c) is relevant with flow resistance58. It is proportional to the inverse square of the width of the microfluidic channel, so it leads to an increase in injected volume as the channel width increases (~ 15 nL for 12 s under 200 µm width; Fig. 4h). Besides, when micropump is actuated repeatedly, it can cause the undesired backflow due to the structural limitation, resulting in reduced efficiency of the micropump. To minimize this limitation, nozzle/diffuser structure is adopted for the air pathway to ensure a one-way direction of net flow adjusted by asymmetric flow design (Supplementary Fig. 10d). The net flow induced by a single micropump increases as the tapered angle increases (maximum flow rate of 8 nL/s; Fig. 4i), but it begins to decrease gradually over 7° because the effect of nozzle/diffuser structure is reduced due to the flow separation of the boundary layer that occurs as the angle becomes wider. The FEA clearly visualizes the difference of vector field of fluid flow under pumping operation with and without nozzle/diffuser structure (Fig. 4j). The maximum flow velocity caused by nozzle/diffuser structure is about 550 µm/s, which enables the control of flow direction with preventing unwanted backflow. Additionally, the input voltage to the microheater, which activates Joule heating, can adjust the flow rate due to the difference in volume expansion at each actuation chamber. The proposed pumping mechanism with engineering parameters adjustment allows precise flow rate control. Under a switching frequency of 0.75 Hz controlled by a LabVIEW software (Supplementary Fig. 11,12), the flow rate of the drug reaches up to 200 nL/min with an applied voltage of 7 V. The flow rate with ~ 7 nL/min is observed at an applied voltage of 4 V, as shown in the experimental data (Supplementary Fig. 13,14).
The FEA shown in Fig. 4k illustrates a programmable digital sequence for microheater actuation and associated multimodal changes such as the temperature of the actuation chamber, air pressure applied to the reservoir and injection volume of the drug during the peristaltic pumping process with optimized parameters (R, 0.8 mm; W, 100 µm; θ, 10°). With an operating frequency of 0.75 Hz and applied voltage of 4 V, localized temperature oscillates by Joule heating according to the digital input sequence. Per each pumping cycle, when all microheaters turn on by Joule heating (step 1), induced positive air pressure starts to release the drug. Continuous drug injection is allowed by the cyclic operation of temperature oscillation. As the cycle progresses, temperature oscillation stabilizes due to the residual temperature, which means that the applied pressure per cycle becomes saturated. Compared to commercial drug infusion module with hundreds of nL/min for flow rate, proposed microfluidic probe provides suitable flow rate for pharmaceutical stimulation because flow rate ranges from a few nL/min to several tens of nL/min (Supplementary Fig. 15). The proof-of-principle experiment of drug release with dyed fluid into solution (water) for 4 hours proves that the drug delivery system with peristaltic pumping mechanism can make reliable, continuous drug infusion, which is useful to realize artificial homeostasis as the neurochemical stimulation (Supplementary Fig. 16).
In vivo, real-time monitoring of the striatal DA dynamics in normal and PD mice model. For in vivo, real-time monitoring of the striatal DA dynamics with electrical stimulation (Fig. 5a), the MDD DA-sensing probe and the electrical stimulation electrode are implanted in the striatum (Str) and the medial forebrain bundle, respectively (Supplementary Fig. 17a). As illustrated in Fig. 5b (right), the MFB of the nigrostriatal pathway links the substantia nigra pars compacta (SNpc) and the CPu of the Str. We stimulate the MFB electrically with an amplitude of 200 µA and a frequency of 60 Hz to trigger the DA release in the CPu (Supplementary Fig. 17b). In parallel, we measured amperometric responses of the MDD DA-sensing probe to the DA dynamics incurred by the electrical MFB stimulation. The DA signal in the Str was gradually increased during the electrical MFB stimulation (blue line, yellow region), whereas the DA signal was restored to the normal level after cessation of electrical stimulation, likely due to the homeostatic recovery (blue line, blue region)60. In contrast, without electrical stimulation applied to the MFB (control, green line), we didn’t observe any significant change in the DA signal (~-46 nA). The DA dynamics after electrical MFB stimulation were consistent in all animals tested (n = 3, Supplementary Fig. 17c-e). The average current change was measured at about 2.8 nA (blue bar) with MFB stimulation, but remained baseline without electrical stimulation (red bar) (Supplementary Fig. 7f).
Moreover, we tested the biocompatibility of the MDD DA-sensing probe with chronic implantation in the brain. Using immunohistological analysis of reactive glial markers, we assessed neuroinflammatory responses to chronic implantation of the DA-sensing probe. Astrocytic deposition and microglial activation were monitored with immunostaining of an astrocyte marker Gfap (red) and reactive microglia marker Iba1 (green), respectively. No observation shown in any significant difference in neuroglial inflammatory responses between the DA probe-implanted group (w. implantation) and the sham-operated group (w/o implantation) (Supplementary Fig. 18). Collectively, these data demonstrated the functionality of the MDD DA-sensing probe to monitor DA dynamics in vivo, as well as its biocompatibility with brain implantation.
Next, we applied the MDD DA-sensing probe to monitor the DA dynamics in pathological in vivo (Fig. 5d). The DA level was depleted and further replenished with pharmacological treatment. To this end, we adopted a hemi-parkinsonian (hemi-PD) mouse model in which dopaminergic neurons have degenerated only in the lesioned hemisphere61. With this hemi-PD model, we could easily monitor the differences in DA level at CPu with degenerated dopaminergic neurons (right hemisphere) and normal region (left hemisphere), and the dynamic changes in DA level before and after pharmacological treatment of L-DOPA, a DA precursor, because of its DA-depletion status. To generate the hemi-PD model, we injected 6-OHDA, a dopaminergic neuron toxin into the unilateral MFB prior to behavioural and histological validation two weeks post-surgery. In the locomotive behaviour test, the hemi-PD model mice displayed ipsiversive rotation toward the lesioned hemisphere. But after apomorphine (APO) injection, their rotational behaviours were dramatically switched from ipsiversive to contraversive direction (Fig. 5e,f). After the behavioural test, all the mice were sacrificed and perfused for histological validation of unilateral degeneration of dopaminergic neurons. STR and SNpc regions were immunostained for tyrosine hydroxylase (TH, Green), a dopaminergic neuronal marker and dopamine transporter (DAT, Red), a dopaminergic axon terminus marker. As compared to the intact hemisphere (left), the 6-OHDA lesioned side (right) showed complete loss of TH and DAT in the Str and also TH in the SNpc, suggesting unilateral degeneration of dopaminergic axons at the CPu and their somas at the SNpc (Fig. 5g and Supplementary Fig. 19).
After these behavioural and histological validation, we could monitor the differences in DA level at CPu with degenerated dopaminergic neurons (right hemisphere) and normal region (left hemisphere) by implanting the MDD DA-sensing probes in the striatum at both hemispheres of the hemi-PD mouse model’s brain as depicted in Fig. 5h and Supplementary Fig. 20. As shown in Fig. 5i, the largest DA signal (~ -42 nA) were observed at the normal region control of the brain (left hemisphere). On the other hand, in the DA depletion region due to the degeneration of dopaminergic axons, very low electrochemical signals (~ -0.01 nA) were obtained with the MDD DA-sensing probe. Moreover, we applied hemi-PD mice model to further validate the functionality of the MDD DA-sensing probe during pharmacological synthesis of DA from its precursor L-DOPA. The MDD DA-sensing probe was implanted into the lesioned Str in the hemi-PD mice model under anesthesia status (Fig. 5j and Supplementary Fig. 21). Interestingly, DA signal in the MDD DA-sensing probe was gradually increased after L-DOPA injection, as compared to no injection control. About 10 min post injection, the DA level was saturated to the highest level and the high level of DA was sustained for as long as about 1 hour (Fig. 5k). Collectively, these data demonstrated that MDD DA-sensing probe is compatible with real-time measurement of pharmacological DA dynamics in the Parkinson’s disease (PD) mice model.
Recent advances in the treatment of PD have included the development of therapeutic strategies10. Oral administration of L-DOPA is conventionally used as pharmacological treatment11,12, but this approach still has vital side effects such as excessive oscillation of L-DOPA, motor fluctuation and dyskinesia13,14. Compared to conventional oral administration, closed-loop neuromodulation in the electroceutical approach actively regulates the homeostasis in the brain by performing acute treatment based on real-time dopamine concentration. Recently reported electroceutical devices for electrophysiology signals sensing and electrical stimulation have been used for closed-loop neuromodulation, but exhibit the drawbacks of vulnerability to artefacts from the nearby stimulation electrode and frequent adjustment of stimulation parameters31,42. On the other hand, direct monitoring of neurochemicals allows real-time characterization of neural activity with high spatial resolution, signal specificity and interference-free detection from other electrophysiological signals in comparison with closed-loop DBS with electrical neural recordings. In order to implement the artificial homeostasis system by mimicking the brain system, we have reported an implantable MDD DA-sensing probe and a microfluidic probe that have the potentials to provide synchronized neurochemical monitoring and pharmaceutical treatment on the targeted spot for neurodegenerative diseases. The proposed single-unit probe with an all-in-one structure that integrates three essential electrodes allows minimally invasive insertion. Through the enzyme immobilization on ZnO NRs, the surface area of WE is maximized, which enables the DA-sensing probe with a sensitivity of 1 µM DA to detect low DA levels in the brain. The FEA for the mechanical stress shows mechanical reliability of a double-side DA-sensing probe with serpentine electrodes under possible deformations in the brain by reducing the induced stress on the electrodes by almost 20 folds. Additionally, the microfluidic probe with a peristaltic pump enables sustainable and controllable drug delivery with a flow rate of 7 ~ 200 nL/min for chronic pharmaceutical treatment. Theoretical and experimental analysis for design parameters and applied voltage support adjustable flow rate. We have demonstrated the feasibility of long-term drug injection for 4 hours and real-time monitoring of DA levels in vivo. The proposed probes could serve as potential tools for building artificial homeostasis by a bio-responsive closed-loop system with an electropharmaceutical approach. This novel platform and successful demonstration with the hemi-PD model provide new opportunities for chemical closed-loop treatments for extensive neurodegenerative diseases that have been restricted with traditional electroceutical approaches.
Preparation of substrate of the multi-deformable double-sided DA-sensing probe. Spin-casting photocurable epoxy (SU-8 2100; Kayaku Advanced Materials, MA, USA) to develop mould for substrate layer. Spin-casting the polydimethylsiloxane (PDMS; Dow Corning, Sylgard 184) onto the developed mould. Taking-off the substrate layer for the MDD DA-sensing probe with a tweezer carefully.
Fabrication of top component (WE/RE) of the multi-deformable double-sided DA-sensing probe. Spin-casting poly(methyl methacrylate) (PMMA; ~ 1 µm in thickness, Microchem, USA) formed a thin sacrificial layer on a glass substrate (55 x 72 x 1 mm3) to facilitate release. Spin-casting and performing photolithography of a photocurable epoxy (SU-8 2005; 5 µm in thickness, Kayaku Advanced Materials, MA, USA) yielded an overcoat on the PMMA, depositing thin layers of metal by electron beam evaporation (Cr/Au; 7 nm/200 nm in thickness), performing photolithography, wet-etching to create the WE and RE of MDD DA-sensing probe. Depositing thin layers of ZnO (25 nm in thickness) and yielding on the WE surface as a seed layer by RF sputtering and lift-off-techniques. Spin-casting and photolithography of another layer of a photocurable epoxy, followed by spin-casting and photolithography of the masking layer. Oxygen reactive ion etching (100 mTorr, 20 sccm, O2, 150 W, 20–30 min) of multiple polymer layers (epoxy-epoxy-PMMA). Spin-casting photocurable epoxy to form a masking layer and opening only on WE surface by photolithography for a hydrothermal process (3 h at 75℃ of reaction nutrient solution) for ZnO NRs growth with reaction nutrient solution composing of a 1:1 ratio of zinc nitrate hexahydrate (Sigma-Aldrich, USA) and hexamethylenetetramine (HMTA) (Samchun, Korea). Removing masking layer with acetone thoroughly and bake grown ZnO NRs (15 min at 150℃ on a hotplate). Performing ultra-sonication (1 min) to remove residues. Dissolving the PMMA by immersion in acetone to allow release with water-soluble tape (Water-Soluble Wave Solder 5414, 3M, USA). Deposition of Ti/SiO2 (50 nm/200 nm in thickness) by RF sputtering at the back side of the water-soluble tape as a bonding layer for subsequent integration with the substrate of the MDD DA-sensing probe.
Fabrication of bottom layer (CE) of the multi-deformable double-sided DA-sensing probe. Processes similar to those described in the previous section except for deposition and patterning thin layers of metal (Ti/Pt; 20 nm/100 nm in thickness) used for the counter electrode in which deposited by electron beam evaporation and patterned by lift-off-techniques without any additional layers or materials.
Fabrication process for double-sided form of DA-sensing probe. Oxygen plasma-treatment treatment (100 W/20 mtorr/32 s/22 sccm O2 in O2 plasma system, Femto Science, Korea) allows the activation on the surfaces of the substrate, as prepared in the previous section, on both sides, followed by aligning and bonding the prepared top component (WE/RE) and bottom component (CE) on each side of the substrate. After integrating two components with the substrate, heating (70℃ for 30 min) allows strong bonding between the bonding layers (Ti/SiO2; 50 nm/200 nm in thickness) and the substrate. Dissolving the tape by immersing it in DI water is the end of the fabrication process of the double-sided form of the DA-sensing probe, followed by printing Ag/AgCl paste (Sun Chemical, USA) onto the RE and drying the Ag/AgCl paste in the oven (70℃ for 2 h).
Enzyme immobilization process for multi-deformable double-sided DA-sensing probe. Dipping MDD DA-sensing probe tip into prepared tetraethoxysilane (TEOS; Sigma-Aldrich, USA) mixture (ethanol/DI/aqueous ammonia solution (Samchun, Korea)/TEOS; 30 ml/2.5 ml/1.5 ml/100 µl) achieved silanisation (-Si-(OH)2) of ZnO NRs (2 h under stirring condition at room temperature). After the silanisation process, washing the tip of the MDD DA-sensing probe with ethanol thoroughly several times, followed by drying the remaining liquid on the MDD DA-sensing probe at 60℃ in an oven, removed the residues of silane. Adding 3-aminopropyltriethoxysilane (APTES; 40 µl, Sigma-Aldrich, USA) into previously prepared TEOS mixture developed APTES mixture for amine-functionalization on WE surface of MDD DA-sensing probe. Dipping the tip of the MDD DA-sensing probe into the prepared APTES mixture (4 h under stirring condition at room temperature) functionalized the silanised-ZnO NRs with amine group (-NH2). After the amine-functionalization process, cleaning the tip of the MDD DA-sensing probe by an ultrasonic bath with DI water and ethanol, followed by drying the remaining liquid on the MDD-DA sensing probe, removed residues from the reaction. For the promising enzyme immobilization onto the amine-functionalized ZnO NRS, dipping the tip of the MDD DA-sensing probe (3 h under stirring condition at room temperature) into the glutaraldehyde (GA; 25% water solution, Sigma-Aldrich) solution (2.5% in 0.1M PPB, pH 5.8) as a crosslinking reagent. After crosslinking process, washing the tip of the MDD DA-sensing probe, followed by dipping the tip of the MDD DA-sensing probe into the enzyme (Tyrosinase; EC 1.14.18.1 2840 U mg− 1, from mushroom, Sigma-Aldrich, USA) solution (5 mg/ml of tyrosinase in 0.05 M PBS) for overnight at 4 ℃. After the enzyme immobilization process, thoroughly wash the unlinked enzyme away with DI water.
In vitro and in vivo electrochemical measurement of the multi-deformable double-sided DA-sensing probe. In vitro characterization was performed in a PBS solution (0.05 M, pH 7.4, Sigma-Aldrich, USA) to evaluate and optimize the performance of the MDD DA-sensing probe. The in vitro and in vivo electrochemical measurements were obtained by a potentiostat (Interface 1010E, Gamry Instruments, USA). The CV was performed between the potential of -0.5 V and 0.5 V at a scan rate of 150 mV/s, and the amperometric response was performed at a potential of -300 mV vs. Ag/AgCl. The longevity test of the MDD DA-sensing probe was examined by immersing it in a beaker with 0.05M PBS solution (10 mL) by injecting 10 µl of dopamine hydrochloride (10mM, Sigma-Aldrich, USA). After each test, the MDD DA-sensing probe was stored in the beaker with 0.05M PBS solution until the next experiment. Reagents used for the selectivity test such as ascorbic acid, norepinephrine, and levodopa were purchased at Sigma-Aldrich, USA.
Fabrication of a microfluidic probe with peristaltic pump. Prepare four individual constituent parts: the substrate layer with microheater; the pumping layer; the microfluidic layer; and the drug reservoir as follows. Depositing thin layers of metal by electron beam evaporation (Cr/Au; 7 nm/200 nm in thickness), and performing photolithography followed by wet etching defined serpentine traces on the glass substrate (300 µm in thickness) to prepare the substrate layer with the microheater. Spin-casting photocurable epoxy (SU-8 2150; Kayaku Advanced Materials, USA) on a silicon wafer to prepare mould-defining geometries of the pumping layer to achieve thermos-pneumatic actuation by allowing space for air expansion by the microheaters. To facilitate release of patterned PDMS from the mould, the SU-8 mould was treated with evaporated anti-stiction agent (trichloro(1H,1H,2H,-perfluorooctyl) silane, Sigma-Aldrich, USA) for 2 h at room temperature. Casting degassed 10:1 PDMS (base: curing agent) on the mould, followed by clamping with a glass slide (76 x 52 mm2) that was treated with Pt inhibitor solution (5% AEAPS; 3-(2-Aminoethylamino) Propylmethydimethoxysilane, Sigma-Aldrich, USA) and 95% methanol. The PDMS between mould and the glass slide was cured for an hour at 70°C in the oven, and the glass slide with the patterned pumping layer was delaminated from SU-8 mould. For the microfluidic layer, processes are similar to those described in the previous fabrication processes of pumping layer. Briefly explain, spin-casting photocurable epoxy on a silicon wafer to define micropump channel and microfluidic channel geometries of microfluidic layer, followed by anti-stiction treatment as described above. Casting degassed 10:1 PDMS on the prepared mould and clamping the mould with prepared glass substrate treated with Pt inhibitor solution, followed by curing PDMS for an hour at 70°C in oven. The glass slide with the patterned microfluidic layer was detached from the mould. A puncher (1.2 mm ID) was used to make holes at the outlet of micropump channel and inlet microfluidic channel. Spin-casting and curing of flat, thin 10:1 PDMS layer prepared on a PI film (125 µm in thickness) to bond with patterned microfluidic layer to form microfluidic probe and pumping membrane via oxygen plasma treatment (100 W/20 mtorr/32 s/22 sccm O2 in O2 plasma system, Femto Science, Korea) which allow strong bonding between microfluidic layer and prepared thin PDMS by activating their surfaces. For drug reservoir development, fabricating mould for reservoir (42 x 46 mm2) consisting of a 4 x 4 array of chamber structures by 3D printer (Project 3500; 3D Systems, USA). Casting degassed 5:1 PDMS (base: curing agent) onto the 3D printed mould, followed by curing at 70°C in oven for about an hour. After completing the fabrication of the four individual parts for microfluidic probe, strong siloxane bonding via oxygen plasma treatment (100 W/20 mtorr/32 s/22 sccm O2) allows integration between three constituent layers: substrate with microheater; micropump layer; microfluidic layer. The patterned drug reservoir was attached to the top part of microfluidic layer where covering the outlet of the micropump channel and inlet of microfluidic channel using silicone adhesive to avoid any leakages.
Finite element analysis of peristaltic pumping mechanics in microfluidic probe
The injection volume and the pressure in the micropump were calculated using numerical simulation tool (COMSOL). As shown in Fig. 4a, the micropump consisted of heater, pump, and microchannel. For FEA analysis of pumping mechanics, it is necessary to calculate not only the individual working of all components, but also their interactions. Sequentially applied voltage to the heater (Cr/Au; 7 nm/200 nm in thickness) induced temperature fluctuations of each heater due to Joule heating. There was convective heat flux between the external air (20℃) and the device, which decrease the temperature of the device. The heat generated by the microheater is conducted to the air inside the cavities. The air pressure inside the cavity was applied to the walls of the inner boundary containing the peristaltic membrane. The pressure was calculated through the ideal gas equation. The fluid flow and pressure inside the microchannel were calculated using the Fluid-Structure Interaction (FSI) module, and a moving mesh was used to analyse the fluid motion. The meshes on the boundary of the microchannel were formed with extremely high density for calculation accuracy. The material parameters (heat capacity at constant pressure, density, thermal conductivity, electrical conductivity, Young’s modulus, Poisson ratio) used in the FEA were listed in table S1.
Finite element analysis of soft mechanics of multi-deformable double-sided DA-sensing probe. 3D solid mechanics and layered shell modules were used to predict the stress distribution of the MDD DA-sensing probe for various deformation modes. The device consisted of three components: top component (SU8/Cr/Au/SU8; in thickness 5 µm/ 7 nm/ 200 nm/ 5 µm); bottom component (SU8/Ti/Pt/SU8; in thickness 5 µm/7 nm/50 nm/5 µm); and substrate (PDMS 250 µm). The other components (enzyme, ZnO, Ti/SiO2 and Ag/AgCl) were ignored in the simulation since they had negligible effect on the stress. The simulation of the PDMS substrate exploited the hyperelastic model (Money-Rivlin model; C10 = 0.06 MPa; C01 = 0.046 MPa). The stress of the MDD DA-sensing probe was analysed by applying torque and force suitable for each mode. The material parameters are listed in table S1.
In vitro biocompatibility test. To validate biocompatibility of the encapsulating MDD DA-sensing probe materials, SH-SY5Y neuroblastoma cells (CRL-2266, ATCC, USA) were used for live/dead cell assay. SH-SY5Y cells were supplemented with Dulbecco’s Modified Eagle’s Medium (DMEM) containing 10% fetal bovine serum (FBS) and 1% penicillin/streptomycin. SH-SY5Y cells were plated at a population of 105 cells/ml in a petri dish (D7804-500EA, Nunc, USA) with a DA-sensing probe fixed to the bottom and incubated at 37℃ under a humidified 5% CO2 atmosphere. After incubation for a predetermined time (3 days), the medium was removed, and the MDD DA-sensing probe was washed in Dulbecco’s phosphate buffered saline (DPBS). Live and dead SH-SY5Y cells were stained by incubating in assay reagents (calcian AM for labelling live cells and ethidium homodimer-1 for labelling dead cells) containing DPBS for 30 min. The live and dead cell images were taken by a confocal microscope (LSM 700, ZEISS, Germany).
Electrical stimulation in medial forebrain bundle (MFB). For in vivo validation of the MDD DA-sensing probe, we induced enhancement of dopamine levels in the striatum by MFB stimulation. To prevent collision of the two stereotaxic arms, the stimulation electrode (NeuroNexus, MI, USA) was inserted at the left MFB with an angle of 20° posteriorly at the following coordinates: A/P = -4.4; M/L = -1.2; D/V = -7.8. The electrical stimulation consisted of biphasic pulses (2 ms per pulse) with an amplitude of 200 µA and a frequency of 60 Hz.
In vivobiocompatibility test. To confirm the biocompatibility in vivo, we implanted the probes in the dorsal striatum of the mouse at the following coordinates: A/P = + 1.0; M/L = -1.5; D/V = -2.75. After 3 weeks of recovery, mice were anaesthetised with tribromoethanol solution delivered by intraperitoneal injection (250 mg/kg), and perfused with phosphate-buffered saline (PBS, pH 7.4) followed by 4% paraformaldehyde (PFA). To examine an inflammatory response after probe implantation, immunohistological analysis for glial cell markers was performed. Sliced brain tissues were preincubated with blocking buffer (2% normal goat serum, 0.2% bovine serum albumin, 0.3% Triton-X 100 in PBS) for 1 h at RT, and transferred to a primary antibody solution (Anti-Iba1 (Wako, 019-19741), anti-Gfap (Abcam, ab4674) diluted in the blocking buffer) for 24 h at 4°C. And tissues were washed with washing buffer (0.2% Triton-X 100 in PBS) 15 min, 3 times. After washing, the brain sections were incubated in a secondary antibody solution containing DRAQ5 (ThermoFisher, 62254), Goat anti-Rabbit IgG (H + L) conjugated with Alexa Fluor 488 plus (ThermoFisher, A-48282), and Goat anti-Chicken IgG (H + L) conjugated with Alexa Fluor 568 (ThermoFisher, A-11041), for 3 h at RT. And tissues were washed with the washing buffer for 15 min, 2 times, and PBS for 15 min, 1 time. Tissues were mounted on a glass slide, and confocal imaging was performed.
Animals. Adult Sprague-Dawley male rats, 250-350g (Samtako, South Korea), and adult C57BL/6J mice, 25-35g (DBL, South Korea), were used for in vivo experiments. The animals were housed in groups of 2–3 rats per cage and 3–5 mice per cage under a 12h light/dark cycle with free access to food and water ad libitum. All in vivo experiments were approved and conducted following the guidelines of the Institutional Animal Care and Use Committee at Daegu Gyeongbuk Institute of Science and Technology (DGIST-IACUC-21081101-0002).
In vivo dopamine sensing under electrical stimulation. To measure dopamine change using DA-sensing probes in vivo, normal rats were used. The rats were anaesthetized with 1–2% isoflurane and placed in a stereotaxic instrument frame. The DA-sensing probe was tested in vivo through an electrical experiment. MDD DA-sensing probe was inserted in the left stratum at the following coordinates (from bregma): anterior-posterior (A/P) = 1.3; medio-lateral (M/L) = -2.5; dorso-ventral (D/V) = -4.5. For inducing dopamine release, a stimulation electrode (NeuroNexus, MI, USA) was inserted in the medial forebrain bundle (MFB). To prevent collision of the two stereotaxic arms, the electrode was inserted at the left MFB with an angle of 20° posteriorly at the following coordinates: A/P = -4.4; M/L = -1.2; D/V = -7.8. The electrical stimulation consisted of biphasic pulses (2ms per pulse) with an amplitude of 200 µA and a frequency of 60 Hz. All experiments were approved and conducted following the guidelines of the Institutional Animal Care and Use committee at Daegu Gyeongbuk Institute of Science and technology (DGIST-IACUC-21081101-0002).
Unilateral Parkinson’s model generation and its validation. To measure dopamine changes in a pathological condition, we generated a unilateral Parkinson’s mice model using 6-OHDA toxin. 30 min before 6-OHDA infusion, the mice were intraperitoneally injected with pargyline hydrochloride (5 mg/kg) and desipramine hydrochloride (25 mg/kg) solution. The mice were anaesthetized with isoflurane and mounted on a stereotaxic frame. 6-OHDA (Sigma-Aldrich, H-116) solution was delivered into MFB at the following coordinates: A/P = -1.20; M/L = ± 1.10; D/V = -5.00. After the surgery, post-operative care was performed on each mouse. Ibuprofen was administrated orally in drinking water for 7 days, and 5% glucose-saline was injected subcutaneously for rehydration and nutrition for 14 days. After 2 weeks of surgery, mice locomotion was recorded on Phenotyper (Noldus). The basal condition was recorded in 10 min, and apomorphine (Tocris, #2073) was subcutaneously injected and recorded in 10 min. Ethovision XT™ (Noldus) software was used for behavioural analysis. Nose and body point tracking data was used for rotation direction and numerical calculation. All behavioural tests were performed, and the mice were perfused for immunohistochemistry. In this immunohistochemistry experiment, anti-TH (Abcam, ab76442) and anti-DAT (Millipore, MAB369) were used on primary antibodies, and goat anti-Chicken Alexa 488 (ThermoFisher, A-11039), goat anti-Rat Alexa 568 (ThermoFisher, A-11077) were used on secondary antibodies, and DRAQ5 was used for nuclear staining.
In vivo dopamine sensing in Parkinson’s mice model. Unilateral Parkinson’s model mice were anaesthetized with isoflurane and fixed on a stereotaxic frame. The DA-sensing probe was implanted on the dorsal striatum and started probe pre-conditioning for 30 min. After the pre-conditioning, the recording was started. During the current recording, levodopa (10 mg/kg) was intraperitoneally injected into the anaesthetized mice.
Acknowledgements
This work was supported by the National Research Foundation of Korea (NRF) funded by Ministry of Science and ICT (NRF-2017M3A9G8084463, NRF-2019M3C1B8090845, CPS21101-130, MOTIE 20006400, HR22C183201).
Author contributions
H.H.J., H.L., J.H., J.R.P., Y.-S.-O. and K.-I.J. designed the project and wrote manuscript; H.H.J., H.L., J.R.P., S.K., J.K., J.Y., S.O. J.J., S.S., J.S., T.S.Y., Y.L., J.W., K.S.L., J.-W.C., J.-W.J., S.K., J.-C.R., H.K. and Y.-S.O. carried out experiments and analysed the experimental data; J.H. led the fluid and solid mechanics modelling; and Y.-S.O. and K.-I.J. supervised the project. All of co-authors contributed to the analysis of experimental results. The diagrams, including mouse model (Fig. 5b,h,j), were created with BioRender.com.
Competing financial interests
The authors declare no competing interests.
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There is NO Competing Interest.
- Jangsupplementary.pdf
Artificial homeostasis systme with electrochemical sensing and pharmaceutical stimulation for impaired brain function