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A microwell-based impedance sensor on an insertable microneedle for real-time in vivo cytokine detection | Microsystems & Nanoengineering
Results and discussion . The sensor, consisting of a 20?μm?×?20?μm microwell array comprising 25 individual 2?μm-diameter wells embedded on a sensing tip, is lithographically configured on a laser micromachined fused silica microneedle. A schematic view of the microwell sensing platform, comprising a pair of gold electrodes separated by a 40?nm insulating layer of aluminum oxide, is shown in Fig. 1a and b . The basic principle of this impedance sensor is the continuous measurement of the impedance change between the sensor electrodes arising from the specific binding of the target protein to the probe antibody already present within the microwells. The probe antibody was first immobilized on the gold electrode surface inside the microwells, as shown in Fig. 1c . Specific binding of the target protein to this antibody (Fig. 1d ) affects ion transport inside the microwells, resulting in a change in impedance between the two electrodes. As a result, real-time monitoring of both the antibody attachment process and target protein binding can be achieved by continuously capturing this change in impedance. Fig. 1: Microwell-based sensing platform. a A schematic view of the microwell sensor and ( b ) zoomed-in side view of the multilayer sensing structure. Schematic images showing ( c ) the antibody functionalized sensor surface inside microwell, and ( d ) specific binding of the target protein to preimmobilized antibody Full size image The key sensor fabrication processes are schematically illustrated in Fig. 2a . The sensors were prepared on 100?mm diameter, 500-micron thick fused silica substrates. A first adhesion layer of 5?nm of chromium followed by 100?nm of gold was deposited by utilizing physical vapor deposition (Lesker PVD75 E-beam Evaporator) and defined by using photolithography and lift-off, yielding the lower electrode, interconnecting line, and bonding pad. The first 40?nm aluminum oxide (Al 2 O 3 ) insulation layer was deposited by using atomic layer deposition (ALD, Cambridge Nanotech S200 ALD). A second 5?nm adhesion layer of chromium followed by a 100?nm gold electrode partially positioned on top of the first electrode in the sensing region, as well as a separate interconnecting line and bonding pad, was formed in the same fashion, and a second 40?nm Al 2 O 3 layer was deposited to insulate the sensing platform. To pattern microwells (2?μm diameter) in the electrode overlapping region, a laser lithography system (Heidelberg DWL 66+ Laser Writer) was utilized. The microwells were prepared with sequential reactive ion etching (Oxford Cobra inductively coupled plasma etcher) of two layers of Al 2 O 3 and wet etching of gold until the bottom gold electrode was exposed. Gold bonding pads were lithographically defined and exposed for connection and recording. Each sensor was micromachined into a needle shape using an excimer laser (IPG IX-255 excimer laser micromachining system) and singulated from the fused silica substrate, as shown in Fig. 2d . These microneedle sensors consist of a shank, approximately 4?mm in length and 0.9?mm in width, each bearing a two-electrode configuration. The two electrodes have an overlapping region (Fig. 2f ), 20?×?20?μm in dimension, on which an array of 25 individual 2?μm diameter microwells was configured. External electrical connections to the gold bonding pads were facilitated by the use of commercially available connectors (Digi-Key Electronics, Thief River Falls, MN, USA), as shown in Fig. 2b and e . A layer of polydimethylsiloxane (PDMS, Sylgard 184, Dow Corning, 10:1 prepolymer/curing agent) was applied on the microneedle shank and connecting pads to further insulate and seal the gold interconnect traces and bonding pads. Fig. 2: Schematic illustration and images corresponding to steps for fabricating microneedles bearing microwell sensors. a , b Steps for fabricating microwell sensors, ablating and releasing them from the substrate using an excimer laser, and packaging with connectors. Images showing c microscopic image of wafer-level fabricated sensors, ( d ) sensors released from the wafer after laser micromachining, ( e ) an overview of a packaged sensing platform, and ( f ) close view of tip and sensing region with an array comprising 25 individual microwells Full size image Figure 3b shows the magnitude and phase behavior of the electrochemical impedance as a function of frequency for a typical microneedle sensor when immersed in various solution environments. In both cases, a monotonic decrease in the magnitude of impedance as a function of frequency was observed. The measured trend of the phase angle as a function of frequency was more complex. To relate these trends in electrochemical responses to the electrical properties of the microneedle sensor, the EIS data were fitted into the equivalent circuit model shown in Fig. 3a . The circuit components are hereby described with reference to the corresponding physical phenomena that they model. The equivalent circuit model is categorized into two sections indicated with dashed lines in Fig. 3a : outside sensing region and microwell sensing region. Outside sensing region: we modeled the effects of two thin oxide-insulated gold traces near the sensing region as a branch comprising a series arrangement of capacitors, C ox02 and C ox03 , representing the dielectric properties of the protecting Al 2 O 3 layers on the lower and upper gold traces, respectively, and a resistor, R solution01 , quantifying the resistance to the movement of ions through the solution in the outside sensing region. Within the microwell sensing region, the capacitor C ox01 reflects the dielectric property of the Al 2 O 3 insulation layer in the overlap region of two gold electrodes. A parallel combination of constant phase elements, C PE01 and C PE02 , represents the interfacial capacitance. These interfacial capacitances, together with charge transfer resistances R ct01 and R ct02 , are adopted to model the two gold–electrolyte interfaces within the microwells. R well is employed to quantify the resistance to the movement of ions through the solution inside the microwells. C ox04 describes the dielectric property of Al 2 O 3 on the top gold electrode, and R solution02 represents the resistance to the movement of ions through the solution in the region adjacent to the microwell array extending from the microwell to the uppermost Al 2 O 3 layer. Fig. 3: Electrochemical impedance of the microwell sensing platform. a A physical representation of the equivalent circuit model after the immersion of label-free sensor into electrolytes. b A set of representative Bode plots showing when antibody immobilization and target protein detection occurs. Experimental results are shown with symbols, and fitting results are shown with solid lines Full size image To determine the parametric values of these circuit components, the equivalent circuit model described in Fig. 3a (solid lines in Fig. 3b ) was fitted to two sets of measured representative impedance characteristics of the microneedle sensor (symbols in Fig. 3b ). The fitting of measured impedance data was performed utilizing Echem Analyst, proprietary software developed by Gamry (Gamry Instruments, Warminister, PA, USA). The fitting results are represented by solid lines in Fig. 3b , and parametric values corresponding to the circuit components are summarized in Table 1 . Overall, the extracted parametric values are reasonably correlated with estimated values calculated from the sensor geometry and electrical properties of materials, as listed in Table 1 . By performing a perturbation analysis, we found that the goodness of fit was not highly sensitive to the parameter R solution01 , further justifying its removal in the simplified model, as discussed in the Supplementary Information . The exploitation of the model to determine favorable impedance parameters and frequencies for use in maximizing impedance changes of the sensor with binding events is shown in the Supplementary Information . The result of this modeling showed that a suitable parameter to monitor binding event-related impedance changes of this particular sensor geometry is the real part of the sensor impedance at a frequency of 100?kHz. Table 1 Summary of the EIS measurement and parameters fitted with the equivalent circuit model. Full size table To validate the adsorption of anti-hIL8 on the sensor surface, fluorescein-conjugated anti-hIL8 (IC208G, Minneapolis, MN, USA) was employed as an adsorption probe. After incubating the sensor with this anti-hIL8 solution for 10?min, the antibody solution in the fluidic cell was removed, the cell was refilled with PBS to remove excess antibody, and the PBS rinse was removed. Consistent with the impedance monitoring result, Fig. 4b shows that anti-hIL8 antibody molecules were successfully immobilized on the sensor surface inside the microwells. In particular, antibody molecules were observed to adsorb selectively on the gold electrode surface, further illustrating that the upper Al 2 O 3 protecting layer provided effective surface passivation against adsorption of antibody on nonmicrowell regions of the sensor. In addition, we observed that the microwell array was only partially functionalized with anti-hIL8, suggesting that the available binding sites for hIL8 might differ from sensor to sensor, which would result in a different percentage change of Z real for different sensors during the antibody immobilization step. One potential explanation is that due to the nonuniformity of wafer-level fabrication, the areas of exposed gold surface inside microwells for sensors located on different regions of the substrate could differ in specific binding affinity. Fig. 4: Prefunctionalizing sensors with probe antibody. a Real part of the sensor impedance measured at 100?kHz as a function of time during antibody adsorption. Immobilization of anti-hIL8 results in an increase in the real component of sensor impedance. b Optical fluorescence microscopy image of the sensor microwell region after functionalization with fluorescein-conjugated anti-hIL8 (green) Full size image The specific detection ability of the sensor for target antigen was investigated in vitro. Nonsingulated sensors on the fused silica substrate (Fig. 5a ) were prepared with fluidic cells as described above. After the anti-hIL8 immobilization step (Fig. 5b ), the fluidic cells were aspirated and rinsed using PBS. PBS containing no target protein was introduced into the cell as a negative control step. Aside from the initial baseline shift common to all steps, the impedance between the electrodes generally decreased, as shown in Fig. 5c . This was seen with all negative control experiments and was attributed to antibody desorption. Introduction of a 1?nM solution of antigen (hIL8) in PBS into the fluid cell caused the impedance to increase; the observed increase in impedance displayed an exponential time course (Fig. 5d ) similar to that observed during antibody immobilization. The increase in impedance was consistent with the hypothesized reduced ion transport rates (increase in ionic resistivity) inside microwells resulting from specific bonding between the hIL8 antigen and preimmobilized anti-hIL8. Fig. 5: In vitro cytokine detection by microwell devices on fused silica wafers. a In vitro protein detection apparatus for a sensor on a fused silica wafer, comprising PDMS fluidic cell and electrical interconnection. b – d Potentiostat data at 100?kHz with a time interval of 50?s. b During the 10-min anti-hIL8 incubation step, the sensor responded with an increase in Z real . c After the removal of excess anti-hIL8 solution, blank PBS containing no target protein was added to the fluidic cells as a negative control, and the baseline impedance generally decreased. d Specific binding of target protein (hIL8) to probe the antibody (anti-hIL8) resulted in an increase in impedance after the initial baseline shift. e Correlation between sensor output and ELISA-based assessment of hIL8 concentration in endogenous mouse serum samples in vitro. The calibration curve showed a positive correlation between the normalized percentage change in the real component of sensor impedance \((\% \bigtriangleup {{{\mathrm{Z}}}}_{{{{\mathrm{real}}}}})\) and the hIL8 concentration. Error bars (where present) indicate the standard deviation of measurements from different sensors measured at the same concentration ( N ?=?2 except at hIL8?=?1000?ng/mL, where N ?=?3) Full size image To further validate detection specificity, the microneedle sensor performance was characterized for undiluted transgenic mouse serum samples with endogenously expressed hIL8. It has been reported that mouse serum has a highly complex proteomic composition and a wide dynamic range of protein concentrations. In a previous study, more than 12,300 unique peptides that originated from 4567 unique proteins were identified 34 . These were only approximately 16% of all known mouse proteins. Quantitative measurement of the cytokine of interest against the background of these potentially confounding proteins would additionally validate the selectivity of the sensor. Samples from different mice contain different concentrations of hIL8. Sensor preparation and prefunctionalization experiments were carried out as discussed previously. In all experiments, a sample of serum from a healthy nontransgenic mouse (Sigma Aldrich, St. Louis, MO, USA) not expected to produce hIL8 was utilized as a negative control, and this was followed by the introduction of mouse serum-containing hIL8. Mouse serum samples with endogenous hIL8 generated a response similar to that of hIL8 in purified PBS buffer, as shown in Fig. 5d , while normal mouse samples exhibited behavior similar to that of PBS (as shown in Fig. 5c ). Non-IL8 proteins in the mouse serum did not interfere with the ability of the sensor to quantitatively determine the presence of IL8. Due to the nonuniform nature of wafer-level microfabrication (e.g., dry etching process to remove the Al 2 O 3 layers and wet etching of metal thin films), sensors located on different regions of the substrate were observed to behave differently in the detection measurements. For example, sensors with differing exposed gold surface areas for immobilizing anti-hIL8 inside microwells resulted in different percentage changes in Z real when these sensors interacted with solutions containing the same biological concentration level. To ameliorate this effect, the percentage change in Z real during anti-hIL8 functionalization was taken to indicate the sensitivity of a particular sensor and used to calibrate subsequent responses of that particular sensor. In particular, we adopted the calibration concept of normalized percentage change in Z real , which is calculated as \(\frac {\% \bigtriangleup {{{\mathrm{Z}}}}_{{{{{{\mathrm{real}},\,{\mathrm{detection}}}}}}}} {\% \bigtriangleup {{{\mathrm{Z}}}}_{{{{{{\mathrm{real}},\,{\mathrm{antibody}}}}}}}}\) , where \(\% \bigtriangleup {{{\mathrm{Z}}}}_{{\mathrm{real},\,{\mathrm{detection}}}}\) represents the percentage change in the real component of sensor impedance in the antigen detection step, and \(\% \bigtriangleup {{{\mathrm{Z}}}}_{{{{{{\mathrm{real}},\,{\mathrm{antibody}}}}}}}\) represents the percentage change in the real component of sensor impedance during antibody functionalization. After measurement by the sensor, aliquots of the various mouse sera were sent for ELISA, and the correlation between the two measurement methods was assessed. Outlier testing was performed for all quantitative data. Figure 5e reveals a positive correlation between the normalized percentage change in the real component of sensor impedance and the hIL8 antigen concentration level in undiluted mouse serum. An increase in the concentration of hIL8 expedited the transfer of this antigen from the solution to the electrode surface and drove the antibody-antigen binding reaction forward. The calibration curve showed a detection range from 100?ng/mL to 1000?ng/mL, with an R 2 of 0.993. Together, these serum sample experiments demonstrated that the sensor el ectrical measurements can be readily attributed to selective antibody-antigen binding, and they validated the ability of the microwell sensor to discriminate against false-positive signals arising from either electronic noise or nonspecific binding. To qualitatively demonstrate the capability of microwell technology to detect other cytokine proteins with high selectivity, we functionalized the sensor with alternative antibodies, including anti-TNFα IgG (R&D Systems, Minneapolis, MN, USA) and anti-IL6 (R&D Systems, Minneapolis, MN, USA), and recorded the real component of impedance-versus-time measurements when the antibody-functionalized sensors were tested with solutions with and without the target proteins. As expected, when the target protein was present in the solution, Z real increased after incubation for 10?min, whereas in the case of the negative control, the sensors responded with a decrease in Z real . These results illustrated that microwell technology is generalizable for monitoring other cytokines. The mechanical stiffness and functional stability of sensors configured on microneedles were subsequently assessed. Sensors were singulated into a microneedle shape and tested on a skin phantom to determine if the microneedle could effectively and reproducibly deliver sensors to the desired location while maintaining sensing functionality. The microneedle was first functionalized with anti-hIL8, then inserted through the skin phantom layer mimicking the dermis and hypodermis (a mixture of gelatin and agar, as shown in Fig. 6a ). After insertion, the sensor was replaced back into the fluidic test cell. Figure 6b–e shows impedance data at 1?MHz obtained when the antibody-functionalized sensors were tested with PBS both with and without the target antigen. Fig. 6: In vitro cytokine detection by a microwell sensing platform in a needle configuration. a Initial wetting of the sensor by adding PBS to the fluidic cell resulted in a significant drop in impedance. b As antibodies were physically adsorbed onto the electrode surface inside microwells, the sensor responded with an increase in impedance. c Microneedle was inserted through a skin phantom. d After sensor removal from the phantom and placement in the fluidic cell, blank PBS containing no target protein (negative control) was added to the fluidic cell, and the sensor impedance baseline generally decreased. e Addition of the target protein hIL8 resulted in specific binding to the probe antibody (anti-hIL8), resulting in an increase in sensor impedance after the initial baseline shift. All impedance measurements were taken at 1?MHz Full size image The specific cytokine detection functionality of the sensor was further validated with a transgenic mouse model in vivo. A transgenic mouse group (positive mice) expressing hIL8 and a normal mouse group not expected to express hIL8 (negative control) were tested. Mice were anesthetized, and microneedles were inserted into the muscle without any damage, thereby successfully positioning the microwells at the desired location with an insertion depth of 3.5?mm. The sensor impedance was monitored immediately after postinsertion. Sensor-based cytokine concentration measurements were obtained using normalized percentage changes in sensor impedance, as discussed above. After sensor measurement, ELISA was performed on extracted mouse serum to quantify the concentration level of hIL8 in each mouse tested. Eight positive mice allowed us to determine sensor responses as a function of the concentration levels quantified by ELISA. The relationship between sensor readouts and the corresponding concentration levels is shown below in Fig. 7e . Reliable detection of hIL8 was shown for concentrations as low as one picomolar. As seen in Fig. 7e , the sensor has a wide detection range spanning four orders of magnitude, from 62?pg/mL to 538.6?ng/mL, with an R 2 of 0.9711. Sensors in eight control mice showed behaviors similar to those for in vitro PBS control steps, i.e., no response. The comparison between these two types of mice allowed us to not only observe specific binding of the target protein in the positive mouse group but also the rejection of the effects of other endogenous mouse proteins that might be present in the blood of both groups. It is also interesting to note that the control sample of healthy mice had a full complement of endogenously expressed proteins, in addition to any extra cytokines or biomarkers that the invasion of the microneedle might cause to be released locally; however, the microwell sensor was insensitive to all of these potentially confounding factors. It was further noted that the sensitivity of the sensors in vivo experiments (as assessed by the slope of the regression line for comparison with ELISA) was larger than that of the in vitro experiments ( Supplementary Information ). This is potentially due to differences in operating temperatures. All in vitro characterizations were conducted at room temperature (23?°C), while the in vivo study occurred at mouse body temperature (37?°C). A higher temperature might accelerate the transport of the antigen to the sensor surface, as well as the antibody-antigen binding reaction rate, resulting in a larger amount of specific binding inside microwells to impede ion transport with the same IL8 concentration level. The detection range measured in this study was limited by the naturally expressed levels of hIL8 in this set of available transgenic mice. Fig. 7: In vivo hIL8 detection by a microwell sensing platform in transgenic mice. a Packaged microneedle sensors ready for animal studies. b An in vivo experimental setup including laptops and potentiostats for data acquisition and workstations for antibody functionalization and animal operations. c , d A sensor configured on a fused silica microneedle was inserted into the thigh muscle of an anesthetized mouse. e Calibration curve of the linear relationship between the normalized percentage change in the real component of sensor impedance \((\% \bigtriangleup {{{\mathrm{Z}}}}_{{{{\mathrm{real}}}}})\) and hIL8 concentrations ranging from 62?pg/mL to 538.6?ng/mL. Error bars (where present) indicate the standard deviation of measurements from different sensors used at the same concentration ( N ?=?2). e Enlarged view of the linear relationship between \(\% \bigtriangleup {{{\mathrm{Z}}}}_{{{{\mathrm{real}}}}}\) and hIL8 concentration at low concentrations Full size image An increase in IL8 expression is observed in many cancer types, such as brain, breast, cervical, colorectal 35 , gastric, lung, melanoma, ovarian, prostate, renal, oral 36 , and thyroid 37 cancers. As a result, elevated IL8 levels have been identified as a prognostic factor 38 . In addition, it has been proven that a high concentration of IL8 is a sign of myocardial infarction 39 . Wong reported that the IL8 protein level is higher in patients with oral cancer (750?±?236?pg/mL) than in control groups (250?±?130?pg/mL) 40 . A serum IL8 protein value of 600?pg/mL is the threshold for discriminating patients with oral cancer from healthy individuals 41 , 42 . In healthy individuals, the basal IL8 clinical level is 5–10?pg/mL 43 . The present in vivo study confirmed that the detection range of the microneedle sensor for in situ hIL8 detections comfortably include the physiological dynamic range for early diagnosis of some specific types of cancer. .
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