Patent ID: 12188002

DETAILED DESCRIPTION OF THE INVENTION

Although preferred embodiments of the disclosure are explained in detail, it is to be understood that other embodiments are contemplated. Accordingly, it is not intended that the disclosure is limited in its scope to the details of construction and arrangement of components set forth in the following description or illustrated in the drawings. The disclosure is capable of other embodiments and of being practiced or carried out in various ways. Also, in describing the preferred embodiments, specific terminology will be resorted to for the sake of clarity.

It must also be noted that, as used in the specification and the appended claims, the singular forms “a,” “an” and “the” include plural referents unless the context clearly dictates otherwise.

Also, in describing the preferred embodiments, terminology will be resorted to for the sake of clarity. It is intended that each term contemplates its broadest meaning as understood by those skilled in the art and includes all technical equivalents which operate in a similar manner to accomplish a similar purpose.

Ranges can be expressed herein as from “about” or “approximately” one particular value and/or to “about” or “approximately” another particular value. When such a range is expressed, another embodiment includes from the one particular value and/or to the other particular value.

By “comprising” or “including” is meant that at least the named compound, element, particle, or method step is present in the composition or article or method, but does not exclude the presence of other compounds, materials, particles, method steps, even if the other such compounds, material, particles, method steps have the same function as what is named.

It is also to be understood that the mention of one or more method steps does not preclude the presence of additional method steps or intervening method steps between those steps expressly identified. Similarly, it is also to be understood that the mention of one or more components in a device or system does not preclude the presence of additional components or intervening components between those components expressly identified.

As shown inFIGS.1A-1F, the present invention is an engineering solution to the challenges associated with monitoring of large area cell cultures with the innovative use of a sensing system100comprising a multimodal sensing platform110comprising an open-mesh structure and a telemetry unit120. The multimodal sensing platform can be a thin, soft sensor array system that can be deployed over the surfaces210of bag bioreactors200. The sensor array can be fabricated using microfabrication processes along with functionalization methods necessary for measuring different modalities of the cell culture300in the bag bioreactor200, for example, the pH, glucose, and temperature. A serpentine design layout112,114and encapsulation strategy130with silicone-based elastomer can allow the sensor system to achieve specific elasticity and modulus, which are preferential mechanical characteristics for the present thin and soft sensor system to enable it to provide the means to monitoring large area culture qualities through the spatial sensing capabilities, culture compatibility, and scalability.

The telemetry unit120can include miniature integrated circuit (IC) components122directly incorporated with thin-film circuits124, allowing for the real-time, on-board data analysis and wireless data communication. The telemetry unit120can comprise an antenna circuit310, multiplexer320, connections330to the sensing platform110, an amplifier340, a Bluetooth chip350and a voltage regulator360.

The telemetry unit120was based on a circuit design developed for a rigid prototype board, and a fabrication and assembly process was invented to complete a flexible wireless telemetry unit. As shown inFIG.1F, the telemetry unit comprises surface mount chip components necessary for Bluetooth Low Energy, a 2.4 GHz antenna circuit, voltage regulation, multiplexing, and analog-to-digital conversion. Structurally, the unit120comprises a thin-membrane, multi-layer copper/polyimide composite interconnection platform, miniature chip components, and elastomeric encapsulation. The interconnection platform was fabricated and completed thin-film was transferred from the PDMS-coated substrate and transferred to a glass slide, where the surface mount chip components are integrated using reflow soldering technique. Finally, the soldered thin-film structure is encapsulated with a low modulus elastomer (Ecoflex 0300, Smooth-On) for full isolation of the electronics. The fully assembled telemetry unit has flexibility and compliant mechanical properties to bend naturally with the underlying surface.

FIGS.1A-1Fshow an overview of the present system100and its application in a bioreactor200. The thin-film flexible electronic circuit124is formed using microfabrication processes, allowing the circuit's thickness to be only 5 μm. The sensor system is integrated in the cell bag's lower membrane and interacts directly with the cell culture. The smart bioreactor is compatible with an available rocking unit400, hence there is no additional cost involved with re-engineering the accessory equipment.

FIG.2illustrates a fully assembled sensing system100integrated onto the inner surface210of a bag bioreactor.

For the fabrication of the multimodal sensing platform110, aerosol jet printing was used (Optomec Aerosol Jet 200, Optomec), a type of additive manufacturing method superior to inkjet printing owing to its ability to print a wide range of materials choice and ink concentration. A poly(methyl methacrylate)-coated four-inch silicon wafer was used as the sacrificial surface to print the sensor array structure using polyimide (PI-2545, HD Microsystems) ink diluted with N-Methyl-2-pyrrolidone (NMP). Upon curing the printed polyimide pattern in a 250° C. oven for two hours, Ag nanoparticle (Ag40XL, UT Dots) mixed with xylene (m-Xylene, Sigma-Aldrich) was printed and sintered at 240° C. for one hour to form the conductive traces.

The top PI is subsequently printed and cured for electrical isolation. Once the additive steps are completed, the sensor structure is transferred to a thin sheet of elastomer substrate. Finally, the necessary electrochemistry and surface functionalization steps for two exemplary chemical sensor types (glucose, pH) take place to complete the sensor functionalization.FIG.2is a photo of the sensor array and its integration with the bioreactor membrane210.

FIGS.3A-3Eschematically illustrate the sensor arrangement (FIG.3A) and the data processing scheme (FIGS.3B-3E). An exemplary 6×6 sensor array110has three representative kinds of sensors130distributed over 10 cm×10 cm area. Three multiplexers320(for example, two 16:1 and one 4:1) serially address the 36 channels and pass the analog data to an analog-to-digital converter (ADC)340. The Bluetooth enabled programmable-system-on-chip350then wirelessly transmits the data to a connected smart device370. The use of multiplexers320significantly reduces the number of wires needed to address all 36 sensors130. The multiplexing speed can be tuned based on the user's requirement.

FIGS.4-6illustrate fabrication methods and the characteristics of three sensor types (FIGS.4A-4D—pH sensor;FIGS.5A-5D—glucose sensor; andFIGS.6A-6D—temperature sensor). For all three sensors, layers of polyimide and sputter-deposited conductors/electrodes were structured using microfabrication processes, such as spin-coating, sputter deposition, reactive ion etching, and wet etching. For pH and glucose sensors, the platinum electrode is functionalized with iridium oxide and glucose oxidase, respectively. For temperature sensing, Texas Instruments' LMT70 chip is soldered directly onto the flexible circuit platform. The characteristics of the three sensor types are shown in the right-most figures.

FIGS.4A-4Bare optical micrographs showing the Pt electrodes before and after IrOxdeposition.FIGS.4C-4Dillustrate the resulting pH sensor exhibited a linear, super-Nernstian response with fast response time.

In an exemplary embodiment, for pH sensing, an electrochemically deposited an iridium oxide (IrOx) film was used for its wide pH response range, fast response time, and high pH sensitivity. The IrOxdeposition solution was dispensed over the Pt electrodes to form a puddle, and a platinized titanium mesh electrode was brought to contact the top surface of the solution. A galvanostatic mode was applied using a power supply with 0.01 A, 1.0 V for 40 minutes. The resulting light-blue IrOxfilm exhibited the expected linear, super-Nernstian response (−76.6 mV/pH) when submerged in three buffer solutions with pH levels of 4.01, 7, and 10.01, verifying IrOx's excellent pH sensitivity.

FIGS.5A-5Billustrate the chemistry and a cross-section diagram of the glucose sensor structure. A Poly(MPC-co-EHMA) (PMEH) overcoat provides the H2O2-permeable protection and is designed to stabilize the sensor output.FIG.5Dshows sensor response in a buffer solution to increasing glucose concentration.

In an exemplary embodiment, for glucose sensing, glucose oxidase (GOD) enzyme was employed and its production of hydrogen peroxide (H2O2) in the presence of glucose and oxygen, where the amperometric response is proportional to H2O02concentration. In order to immobilize GOD to Pt electrodes, PMEH was synthesized and polymerized with 2-methacryloyloxyethyl phosphorylcholine (MPC) and 2-ethylhexyl methacrylate (EHMA), and it was used as a hydrogen permeable protection membrane for GOD. The active sensing material was prepared by mixing 5 mg of GOD and 10 μl of PMEH (10 wt % in ethanol) then applying it over the Pt electrode, followed by curing at 4° C. for three hours. To prevent enzyme leakage, PMEH solution was drop-casted over the sensor and cured at 4° C. for three hours.

To verify the functionality of the sensor, the sensor was submerged in a phosphate buffer solution and added 0.01 g of glucose. A commercial glucose sensor (GluCell® Glucose Monitoring System) was used to monitor the actual glucose concentration throughout the test. The sensor exhibited a transient response to the added glucose. For instance, while the potential increase of −2 mV could be detected from the initial addition of glucose, no meaningful sensor response could be measured from the second addition and on. Consequently, it is suspected that, despite the presence of PMEH as the immobilization enhancer as well as the protection layer, enzyme loss has occurred. Currently, the PMEH curing process is being optimized and the effect of PMEH curing to sensor's stability being validated.

In an exemplary embodiment, for temperature sensing, an analog temperature sensor was used in a miniature surface mount chip package (LMT70, Texas Instruments). The temperature sensor, along with a capacitor and a resistor, was integrated with a thin, flexible interconnection platform using reflow soldering. The flexible temperature sensor was submerged in a water bath for functional verification and its temperature reading was compared to its evaluation module (LMT70, Texas Instruments). As shown inFIG.6D, the thin-film sensor's reading correlated well with that of the rigid PCB counterpart with the slight offset of ˜0.1° C. between the two data. Overall, the fabricated sensor exhibited the sufficient sensitivity to the temperature fluctuation created with a heated water bath, demonstrating its capability as a temperature sensor for the smart bioreactor.

FIG.6Ais a circuit layout in the flexible substrate.FIG.6Bis a photograph of a single flexible temperature sensor with a fan-out pads for wire connection. The zoom-in images ofFIG.6Cshow the result of reflow-soldering chip components.FIG.6Dillustrate a heated water bath test of the thin-film and the evaluation module show that both sensors responded to temperature variation with high correlation.

The innovative sensing system100having the multimodal sensing platform110is capable of monitoring cell quality in a large culture area. The thin, soft electronic structure allows the seamless integration with a bioreactor's membrane while the sensor array captures real-time spatial information of the cells with three sensor types, for example, pH, glucose, and temperature. The present monitoring method provides manufacturers with a type of culture information that was previously not available in conventional system, such as the spatial distribution of cell population and culture areas with non-desirable growth rates or cell state. The use of the sensor system can establish the new standards of large-scale cell manufacturing with increased yield and reproducibility.

The present fully integrated wireless sensing system has been implemented.FIGS.2,3Adepicts such a system comprising the wireless telemetry unit120and the exemplary 6×6 sensor array110. The two-dimensional distribution of the three sensor types is schematically illustrated inFIG.3A. This is only an exemplary embodiment, and those of skill in the art understand that alternative distribution patterns and sensor densities have other beneficial culture qualities.

Optical microscope images shown inFIGS.7A-7Cillustrate the visual appearances of the three types of thin-film solid-state sensors used to monitor pH, glucose, and temperature. In an exemplary embodiment, all three types of sensors are manufactured using standard microfabrication processes along with additional electrochemical deposition steps for pH and glucose sensors for deposition of iridium oxide (IrOx) and palladium iron (PdFe), respectively.

To demonstrate the capabilities of the sensing system toward continuous and wireless cell monitoring, a polystyrene Petri dish is integrated with the sensors followed by introduction of the medium containing muscle stem cells (MSC). Representative pH and glucose data collected during MSC growth exhibited good results.

Accurate, sensitive, and stable monitoring of various culture conditions require that the embedded sensors of the present invention exhibit consistent sensor-to-sensor characteristics when integrated over the bioreactor's membrane. Prior to the inventive techniques herein, manufacturing strategies were found lacking and not able to produce highly consistent thin-film, solid-state chemical sensors with intention to be transferred to flexible substrate.

The present invention further comprises novel microfabrication methods allowing wafer-scale manufacturing of solid-state pH and glucose sensors. It enables wafer-level electrodeposition of flexible solid-state chemical sensors for integration with the bag-embedded conductive traces.

As shownFIG.8A, a four-inch polydimethylsiloxane (PDMS)-coated silicon wafer was used to pattern gold (or platinum) electrodes in the 10×10 array. The large region surrounding the electrode array is used to bond a wire. As shown inFIG.8B, a closer inspection of an individual sensor design reveal that each ‘pixel’ contains the circular sensor area, exposed metal tab for the electrical connection with the bag-embedded interconnection, and four thin bridging traces to electrically connect adjacent pixels. To facilitate the tear-off of the selected sensor without affecting other sensors, the contour of the conductive bridges exhibits the shallowed polyimide (PI)'s width in the middle. The locations of such ‘tear points’ are also shown. (FIGS.8C-D).

The electrochemical deposition of IrOxutilized a three-electrode configuration as shown inFIG.8E, where the pulsed voltage (˜700 mV squares every one second) is applied across the working electrode (WE), for example, exposed gold electrodes, and the counter electrode (CE), for example, a platinum (Pt)-coated wafer. For PdFe deposition, a negative voltage (˜1 V) is applied across the WE and the CE.

For both deposition processes, an Ag/AgCl reference electrode (RE) was used to maintain precise applied voltages. In order to prevent unwanted deposition on the connection pads, a stop-off lacquer was applied over the pads prior to deposition steps (Before/After,FIG.8E).FIGS.8F-Gshow the completed pH and glucose sensors ready to be transferred to the bioreactor.

As noted, to facilitate the tear-off of the selected sensor without affecting other sensors, the contour of the conductive bridges exhibits the shallowed PI's width in the middle (FIG.8C), and the effectiveness of this unique design feature is now described. The 10×10 sensor arrays have been prepared at wafer-scale and the PDMS-coated wafer as well as the tear-off feature allow individual sensors to be removed from the donor wafer and integrated with the bioreactor. This assembly concept is illustrated inFIG.9A.FIG.9Bcaptures the moment the completed sensor pixel is peeled from the wafer using pointed tweezers. The weak adhesion between PI and PDMS allows the sensor to be released effortlessly, whereas the tear points guide the controlled fracture of the bridges.FIG.9Cschematically depicts the method with which the transferred sensor is electrically connected to the printed interconnection embedded in the bioreactor. Ag paint is applied over the connection pads on both the sensor and the interconnection (FIG.9D) and allowed to dry for 30 minutes followed by an elastomer coating to expose only the sensor material to the culture medium.

FIG.10Ashows an exemplary temperature sensor circuitry, which operates via the Wheatstone bridge topology. The fabricated Pt thermistor is connected in the place of Rx in the Wheatstone bridge (Block2). Block1is an input voltage buffer that provides voltage when triggered by a power bus from the microcontroller. Block2uses the principles of resistive voltage division and the exponential dependence of temperature on resistance to calculate the temperature. The output voltage is by definition:

Vo⁢u⁢t=Vin⁡(R⁢⁢2R⁢⁢2+R⁢⁢3+R⁢xR⁢1+R⁢x)(Equation⁢⁢1)

Because Rx is the only unknown, it can be calculated from Equation 1. Once Rx is known, the temperature can be determined by the following relation:

ln⁢⁢(R⁢xR)=β⁢⁢(1T-1T0)(Equation⁢⁢2)

Block3is a differential amplifier that calculates the Wheatstone bridge output voltage.

FIG.10Bdescribes a simplified version of the pH sensor circuit. The pH sensor includes a working and reference electrode, which are designed to produce a differential voltage proportional to the solution's pH. Because the reference electrode is non-conductive, it is referenced to ground, precluding the implementation of a differential amplifier. Instead, the voltage at the working electrode is measured with respect to the reference electrode by a non-inverting, low pass amplifier, with the gain set as the ratio

A⁢v=(1+R⁢1R⁢2).
The voltage transmitted to the Arduino is thus:

Vo⁢u⁢t=(WE-RE)⁢(1+R⁢⁢1∥(-jw⁢c)R⁢2)(Equation⁢⁢3)

FIG.10Cdepicts the amperometric circuit designed for the glucose sensor. The circuit is an improvement upon simple amperometric methods because it allows for real time monitoring of the applied cell potential and uses a load independent voltage buffer to stabilize the input voltage. The overall function of the circuit is to supply a differential voltage between the reference and working electrode in order to facilitate glucose ion transduction on the counter electrode.

Block1is a summing amplifier that receives a voltage from a power bus and sets the output line to the inverted sum of the input voltage and the reference electrode voltage. Block2inverts this voltage back to positive polarity. Block3stabilizes the voltage to be fed into the working electrode. Block4stabilizes the reference electrode voltage before being fed into Block1. Finally, Block5transduces the output current between Counter and Ground into a voltage that is provided to the microcontroller. This voltage is calculated from Ohm's Law given the value of the transduction resistor.

The overall structure of the smart bioreactor sensing circuitry is shown inFIG.10D. Each sensor in the sensor array is interfaced with a unique signal transduction circuit as outlined. Once the signals are converted to raw voltages, the microcontroller in the Bluetooth module will select the signal via multiplexer and sample it with an ADC. After sampling, the digital voltage will be sent to a computer running a C application showing a spatial map of the sensor values in real time.

Overall pH sensing capability in cell culture media was improved. Various strategies lead to the surface stability of the sensor, resulting in monitoring subtle pH for seven days. Overall,FIG.11illustrates the enabling technologies for the IrOxsensor material and top membranes.

In prior embodiments, an IrOxpH sensor was fabricated via pulsed electrodeposition. Even though the method successfully enhanced surface conformality of the film electrode by minimizing oxygen evolution, it still has a long-term stability issue. This instability is due to the film's surface hydroxyl status that changes over time in the media and even in air, resulting in the change in the voltage signal.

To increase the surface stability, an applying voltage of the pulsed condition was controlled as shown inFIG.12A. In prior embodiments, it was fixed at a VCN=1.1 V.FIG.12Ashows a linear sweep voltammogram of oxidation reactions at the surface of Au in IrOxdeposition solution. The reaction starts from an oxidation voltage of 700 mV and slowly increases up to near 1.1 V where an unwanted oxygen evolution occurs.

The voltage was changed from 0.7 to 1.1 V and presented the result inFIG.12B. Without voltage optimization, the sensitivity highly fluctuates in a range of −40 to −65 mV/pH. An optimized sensor that used a 900 mV as the VONshowed enhanced stability, while the sensor was stored in air. Considering most commercial glass electrodes and film-type sensors are supposed to be stored in a buffer solution before use, the present sensor can be provided with the circuit in dried status.

The sensor also provides a sensing result measured in a subtle pH range (FIG.12C). Since the cell productivity is highly affected by any small changes of pH, the voltage was measured with different pH buffer solutions of pH 6 to with a 0.1-0.2 discrepancy. The stabilized film electrode showed a super-Nernstian response (−61 mV/pH) and 0.25 pH accuracy.FIG.12Dpresents the voltage change along with its pH value, which changed little for seven days, indicating that the present sensor can work without calibration before use.

In prior embodiments, Nafion was used as a cation-selective membrane. Despite its perm-selectivity, high robustness, and biocompatibility, other cations including K+can go through the membrane in cell culture media which includes various cations (Na+, K+, Zn2+, Fe2+, etc.) from inorganic salts added for the osmotic balance. To minimize this disturbance, an ion-selective membrane (ISM) was adopted for improving selectivity and obtaining accurate result.

The H+-selective membrane cocktail was prepared by mixing 10 wt % hydrogen ionophore I, 89.3 wt % 2-nitrophenyl octyl ether (o-NPOE), and 0.7 wt % potassium tetrakis(4-chlorophenyl) borate. The solution was sonicated for ten minutes and homogenized with a vortex mixer, while the optimized IrOxsurface was being hydrated in pH 7 solution. Afterwards, the volume of 2 μm cocktail of each membrane solution was drop-coated on the IrOxsurface and dried overnight.FIGS.13A-Cshows the voltage result of three samples of each membrane-coated sensor in cell media. The pH difference to the glass pH electrode (FIG.13A) decreased from 0.4 to 0.1 with the ISM (FIG.13C), indicating the decreased interference with other cations.

The inventive pH sensor works as a potentiometric sensor that reads a voltage difference between working and reference electrodes (WEs, REs). Therefore, a high stability is required for the REs even higher than the working electrode, making sure to reduce any signal errors and sensor failure. However, commercially available REs are fragile, bulky, and thus not seamlessly adaptable for the disposable bioreactor.

FIG.14Ashows a photo of the commercial rod RE, and a thin film RE of the present invention that addresses these problems. The present RE film comprises thin AgCl on Ag layer (FIG.11), providing a highly small form factor that occupies a smaller area (FIG.14B). This film configuration comes with a critical issue on its voltage stability due to the unused of a filling solution. Annealing of the Ag film and chemical stabilizing steps of the film structure were adopted, resulting in resolving detachment of the films from the substrate and stable voltage for one week in phosphate-buffered saline (PBS) solution (FIG.14C). In this way, the present film RE film has a very small voltage deviation less than 0.5 mV for 18 days (FIG.14D). To use the present RE for the potentiometric sensor, polyvinyl butyral (PVB) was applied to provide insensitivity to pH change and used in muscle cell media.

The functionality of an anti-biofouling membrane pH sensor in cell media was tested. All the membranes, including Nafion, ISM, and gel, are highly acceptable and biocompatible for the use in cell culture media. pHEMA (poly(2-hydroxyethyl methacrylate) was used for the anti-biofouling effect.FIG.15Ais a photo partially showing a 24-well plate coated with the different membranes coated on the bottom. A 20-μL solution of 0.5 vol % Nafion was drop-coated for 30 minutes and dried in air for one hour after removing residual.

The same coating protocol was followed for the H+-ISM and pHEMA membranes, but they were dried at 80° C. Afterwards, 1 mL C2C12 growth media was dropped with a same number of the cells. Cell viability was measured in a daily basis by dying the cells with a 1:9 vol % PrestoBlue:growth media solution.

Before measurement, the plate solutions were incubated for more than ten minutes at 37° C. The viability results measured at a recommended condition (Fluorescence; excitation 560 nm, emission 590 nm) presented inFIG.15Bshows that all the membranes are compatible with cell culture media.

Even though the Nafion and ISM showed better viability, its deviation is a lot higher than that of the gel-coated surfaces. This could be due to a degradable effect of the cell culture to the selective membranes. Therefore, the present pH electrode coated with the anti-biofouling membrane of a top of the selective layer showed a higher productivity than the IrOxelectrode or the film with one selective membrane (FIG.15C).

Long-term sensing capability has been achieved, as the present invention improves upon the stability, selectivity, and biocompatibility of the pH sensing electrode and RE.FIG.16Ashows the improved stability in voltage signal measured in cell media for seven days. Conventional sensors lose sensing capability in a short period of time due to the unstable surface, but the present sensor showed seven-day stable reading of the voltage. As a result, the sensor was able to read the pH of C2C12 cell media for seven days as shown inFIG.16B.

The results (day 1: pH=8.00, day 7: 7.95) were compared to a bulky glass electrode (day 1: 8.25, day 7: 8.14), indicating that the pH difference was only 0.2 pH.FIG.16Cshows that the present sensor provides a small and thin form factor, which was adopted in cell culture dishes, while the commercial sensor is bulky and used with a filling solution (left side of the sensor). Collectively, the present pH electrode has stable and cell-compatible surfaces to ensure a long-term measurement of pH in cell media.

The present optimized sensor was used with muscle stem cells (MuSC) to perform cell viability measurements. For the experiment, mice cells were seeded in 1 mL growth media using a multi-well cell culture plates with 24 covered wells.FIG.17Ashows the cell plate used for the viability test. First, the plate was prepared by fixing the present IrOxpH electrodes on the bottom of the plate with Ecoflex. The two different membranes were coated on the top of the sensor. The surfaces of the sensor and control plates were coated with laminin mouse protein and collagen I (rat tail) with 0.3:8.3 μl/ml volume ratio in PBS solution. After a 30-minute immersion, the plates were cleaned with PBS and dried in air for 30 minutes. The same number of cells (5 k) were seeded into the wells with 1 mL cell culture media. The growth media is composed of F10 medium containing 20% horse serum, and 1% penicillin/streptomycin. Lastly, a small amount of basic fibroblast growth factor (bFGF; 1 μl/ml) was added in a daily basis.FIG.17Bshows the viability of the MuSC grown for five days in an incubator (temperature: 37° C., oxygen concentration: 20%).

The results showed that the present electrodes rarely affect the cell proliferation at least for three days compared to the control data. Also, the voltage signal of the present sensor in the MuSC media showed a stable reading, which was calculated as near pH=7 (FIG.17C). The pH measured with present sensor was slightly lower than commercial sensor (pH=7.7), which may be due to the coating layer on the top of the sensor before loading the MuSC.

The present invention further comprises improved functionality of an inorganic glucose sensor. A controlled voltage of the sensing voltage of a PdFe glucose sensor was examined.

In prior embodiments, palladium iron (PdFe) inorganic film was used a glucose sensor. The film electrode deposited by using cyclic voltammetry showed a high sensitivity to glucose in a concentration range of 1-55 mM that covers a normal low and high glucose levels used for culturing cells. However, since the film monitored glucose levels based on an amperometric sensing mechanism, the signal appears as a current differential while the circuit applies a certain voltage for detecting glucose levels. Unlike the potentiometric sensors, the voltage should be as low as possible not to interfere with other bio-chemicals in the cell culture solution. In the present invention, focus was on lowering the applying voltage that was optimized before to 0.26 V by characterizing the film as well as by investigating different linearity dependent on the voltage level.

FIG.18Ashows SEM images of the electro-deposited PdFe films with and without coating a membrane on the top. 5 wt % Nafion was used for selectivity and durability of the thin film sensor. As can be seen, the film was covered by the thick membrane, which compensates any surface roughness and may reduce an unwanted oxidation of Fe in the film. Sensing capability of the resulting electrode was investigated with chronoamperometry that applies a fixed voltage for a short period of time (FIG.18B). At −0.01 V, which is a lot lower than the previous condition, the voltages appear different clearly according to the glucose levels, while their response was stabilized in several seconds. Such improved sensing capability was optimized by controlling the voltage levels from −0.01 to 0.5 V as can be seen inFIG.18C.

Collectively, the present sensor showed a linear response at −0.01, −0.1, and −0.2 V with sensitivity from −34 through −119 to −208 μA/log(mM)·cm2. The sensitivity was comparable with the previous result with a lower voltage condition. At the more negative voltage, none of linear response was monitored, indicating there were no oxidation reactions.

It is to be understood that the embodiments and claims disclosed herein are not limited in their application to the details of construction and arrangement of the components set forth in the description and illustrated in the drawings. Rather, the description and the drawings provide examples of the embodiments envisioned. The embodiments and claims disclosed herein are further capable of other embodiments and of being practiced and carried out in various ways. Also, it is to be understood that the phraseology and terminology employed herein are for the purposes of description and should not be regarded as limiting the claims.

Accordingly, those skilled in the art will appreciate that the conception upon which the application and claims are based can be readily utilized as a basis for the design of other structures, methods, and systems for carrying out the several purposes of the embodiments and claims presented in this application. It is important, therefore, that the claims be regarded as including such equivalent constructions.