Incorporating CMOS integrated circuits in the design of affinity-based biosensor systems

A biosensor system incorporating CMOS integrated circuits. In one type of biosensor system, the biosensor system includes a silicon substrate. The biosensor system further includes active devices fabricated on the silicon substrate. Additionally, the biosensor system includes a plurality of metal layers stacked on top of the active devices. Furthermore, the biosensor system includes a passivation layer covering a top metal layer, where the passivation layer includes an opening configured to expose the top metal layer, where the opening is used as a sensing electrode. Additionally, the biosensor system includes a plurality of probes attached to the sensing electrode.

TECHNICAL FIELD

The present invention relates generally to biosensor systems, and more particularly to incorporating CMOS integrated circuits in the design of affinity-based biosensor systems.

BACKGROUND OF THE INVENTION

Affinity-based detection is a fundamental method to identify and measure the abundance of biological and biochemical analytes and is one of the most important analytical methods in biotechnology. Affinity-based detectors (or so-called biosensors in case of detecting biological analytes) take advantage of the selective interaction and binding (affinity) of the target analyte with immobilized capturing probes to specifically capture the target analyte onto a solid surface. A goal of a detection platform is to facilitate specific capturing and ultimately to produce a detectable signal based on the captured analytes. The generated signals correlate with the presence of the target analytes in the sample (e.g., toxins, polymers, hormones, DNA strands, proteins, bacteria, etc.), and hence are used to estimate their abundance.

To create target-specific signals in biosensors, the target analytes in the sample volume first need to collide with the capturing layer, interact and bind to the probes, and ultimately take part in a transduction process (i.e., a physiochemical process which produces certain measurable electrical, mechanical, or optical parameters produced solely by the captured entities). The analyte motion in typical biosensor settings (e.g., aqueous biological mediums) is dominated by diffusion spreading, which from a microscopic point of view is a probabilistic mass-transfer process (i.e., random walk events for a single analyte molecule). Accordingly, the analyte collisions with the probes become probabilistic processes. Moreover, because of the quantum-mechanical nature of chemical bond formation, interactions between probes and analytes, are also probabilistic, adding more uncertainty to the capturing procedure. On top of these two processes which can be considered the biochemical noise of the system, there may also be a detector and a readout circuitry (e.g., optical scanners for fluorescent-based transducers), which likely add additional noise to the already noisy signal.

Besides the inevitable uncertainty associated with the target analyte capturing and detecting, in all practical biosensors, binding of other species to the probes (non-specific binding) is also possible. Non-specific binding (e.g., cross-hybridization in DNA microarrays) is generally less probable than the specific binding when target analytes and the interfering species have the same abundance. Nonetheless, when the concentration of the non-specific species becomes much higher than the target analyte, non-specific bindings (or essentially interference) may dominate the measured signal and hence limit the minimum-detectable-level (MDL). In biosensors, the MDL may be either biochemical noise or interference-limited, while the highest detection level (HDL), is solely a function of capturing probe density and its saturation level.

Due to such impediments, as of today, the accuracy of biosensors systems does not satisfy the stringent requirements of many high-performance biotechnology applications in molecular diagnostics and forensics. In addition, biosensors systems have not successfully made the transition to portable and compact point-of-care devices because their detection platforms still consist of fluidic systems and bulky detectors.

One proposed solution to address the challenges of biosensor systems is to use semiconductor fabrication technologies to build compact, high-performance, and cost-efficient biosensor systems. It is envisioned that such systems (i.e., lab-on-a-chip platforms), include not only the fluidic (macro or micro) systems and sample preparation processes, but also the integrated transducers.

The challenge of designing sample preparation modules in biosensors, to some extent, has been addressed in recent years, particularly in the form of micro-fluidic and automated liquid handling systems; however, the integration of the detector and readout circuitry has not been addressed. One reason why the integration of the detector and readout circuitry has not been addressed is the technical challenge of manufacturing transducers using custom surface and bulk MEMS procedures. Another reason is performance and cost justification of monolithic integration of all components.

In recent years, the idea of employing Complementary Metal-oxide-semiconductor (CMOS) fabrication processes, which are the most robust and widely used fabrication processes in the semiconductor industry, for biosensors has emerged. The rationale behind this, as opposed to using MEMS or other custom processes, is the unmatched yield, cost-efficiency, and the integration capabilities of CMOS processes. While CMOS processes, from the electronic design point of view, offer huge degree of design flexibility and system integration, they are not very flexible in terms of form factor, transducer design and interface integration. Challenges remain in designing biosensors to take advantage of the CMOS fabrication method. The primary design challenge using CMOS technology is the interface design between the assay and integrated chip (IC) which requires additional post-fabrication processes for compatibility in detecting targets (e.g., analytes).

Therefore, there is a need in the art for incorporating the use of CMOS fabrication processes in the design of affinity-based biosensor systems.

BRIEF SUMMARY

In one embodiment of the present invention, a biosensor system comprises a silicon substrate. The biosensor system further comprises active devices fabricated on the silicon substrate. Additionally, the biosensor system comprises a plurality of metal layers stacked on top of the active devices. Furthermore, the biosensor system comprises a passivation layer covering a top metal layer of the plurality of metal layers in order to protect the plurality of metal layers, where the passivation layer comprises an opening configured to expose the top metal layer, where the opening is used as a sensing electrode. Additionally, the biosensor system comprises a plurality of probes attached to the sensing electrode.

DETAILED DESCRIPTION

As discussed in the Background section, biosensors are one of the most important analytical tools in biotechnology today. These detection systems take advantage of the selective interaction and binding of certain biological molecules to identify and detect different analytes such as toxins, hormones, DNA strands, proteins, bacteria, etc. The fundamental advantage of array based biosensors, which compensate for their limited signal-to-noise ratio (“SNR”), is their capability to detect multiple analytes simultaneously. Today, densely packed biosensor arrays (i.e., microarrays) which detect hundreds or even thousands of different analytes are an integral part of biotechnology.

Certain emerging biotechnology applications, such as high-throughput molecular screening and point-of-use (PoU) molecular diagnostics, necessitate biosensor integration, particularly the interfacing of the biochemical part (assay) with the transducer and the readout circuitry. This is mainly due to the stringent requirements of applications which demand compact, cost-efficient, and disposable systems with a high production yield and robust functionality; a goal which silicon-based integrated circuits technology in general, and CMOS processes in particular can provide.

The dominant biosensor and microarray detection modality is visible-range fluorescence spectroscopy using fluorescent labels as the reporters for target analyte molecules. While alternative “label-free” transduction methods (e.g., electrochemical or magnetic) exist today, fluorescent-based detection still remains the most sensitive and robust method, particularly in DNA detection application. The performance advantages of this detection method over other methods originate from the uniqueness of fluorescence phenomenon which makes the generated signals very specific and less susceptible to biological interference.

Referring toFIGS. 1A-D,FIG. 1-Dillustrate a fully integrated fluorescent-based biosensor microarray system100in accordance with an embodiment of the present invention. System100, as shown inFIG. 1A, includes a packaged CMOS integrated circuit101with a sensing area102. In one embodiment, sensing area102has a thickness of approximately 3 millimeters. Details of sensing area102are provided further below in connection withFIG. 3showing the transducers, emission filter, and readout circuitry, including the analog-to-digital converter (ADC).

To visualize the structures of sensing area102for biological significance, a fluorescent image103of a portion of sensing area102is taken as shown inFIG. 1B. Fluorescent imaging techniques, include, but are not limited to electron microscopy, x-ray crystallography, nuclear magnetic resonance spectroscopy and atomic force microscopy. Fluorescent image103illustrates photo-detectors104A-C. Photo-detectors104A-C may collectively or individually be referred to as photo-detectors104or photo-detector104, respectively. System100may include any number of photo-detectors104and the number of photo-detectors104shown inFIG. 1Bis illustrative.

Fluorescent image103further illustrates targets (e.g., analytes)105A-B. Targets105A-B may collectively or individually be referred to as targets105or target105, respectively. System100may include any number of targets105and the number of targets105shown inFIGS. 1C and 1Dis illustrative.

For each target105, a fluorescent label is captured by a DNA capturing probe. For example, capturing probes106A-C are used to capture fluorophores107A-B using a process referred to as fluorescent labeling in connection with target105A. Similarly, capturing probes106D-F are used to capture fluorophore107C in connection with target105B. Capturing probes106A-F may collectively or individually be referred to as capturing probes106or capturing probe106, respectively. Further, fluorophores107A-C may collectively or individually be referred to as fluorophores107or fluorophore107, respectively. System100may include any number of capturing probes106and fluorophores107and the number of capturing probes106and fluorophores107shown inFIGS. 1C and 1Dis illustrative.

While system100is designed and fabricated for DNA microarrays, the achieved specifications are well suited for other biosensor applications.

The foremost challenge in designing fluorescent-based detectors is the proper excitation of labels and the detection of their emitted signal. The photon absorption of the fluorescent label, denoted by A inFIGS. 1A-D, exposed to an incident photon flux, FX, obeys the Beer-Lambert law. For a thin layer of fluorescent labels, the absorption is given by
A=FX[1−e−a0(λ)N]≈FXa0(λ)N,(1)
where N is the surface concentration of labels with extinction coefficient of a0(λ). The total isotropic photon emission, IE, as a function of QY, the fluorescence quantum, is given then by
IE=QYA≈QYFXa0(λ)N(2)

The major function of a fluorescent-based biosensor is to measure N using IEbased on equation (2) in the presence of FX. Although FXgenerally has a slightly different wavelength from IE, it is typically 4-5 orders larger and therefore needs to be blocked during detection.

In one embodiment, a low-power diode pumped solid-state (DPSS) green laser is used with an output wavelength of 532 nm to create FX. To block FXfrom reaching integrated photo-detector104, a multi-layer thin film dielectric long-pass (edge) optical filter (discussed below in connection withFIG. 3) is designed and fabricated. In one embodiment, the optical filter comprises 20 layers of ZnS (n=2.30) and Na3AlF6(n=1.35) with an overall thickness of 2.1 μm. The measured transmittance of the filter is illustrated inFIG. 2. The filter rejects FXby 98 dB at 532 nm, while having approximately 1 dB loss in the optical passband.

In connection with the filter discussed above,FIG. 2is a graph200of the filter transmittance (in dB)201and photodiode quantum efficiency (QE)202versus the wavelength (in nanometers)203in accordance with an embodiment of the present invention. Referring toFIG. 2,FIG. 2illustrates the filter response which rejected (about −100 dB) the light that was below 532 nm (at the fluorescent excitation band) while passing light that was greater than 560 nm. The range of frequencies that are attenuated are referred to as “stopband” as shown inFIG. 2; whereas, the range of frequencies that can pass through the filter without being attenuated are referred to as “passband” as shown inFIG. 2. Further,FIG. 2illustrates that the emitted light of 570 nm wavelength will not be attenuated very much thereby allowing the transducer and detection circuitry (discussed below in connection withFIG. 10) using the CMOS process to detect this light signal.FIG. 2further illustrates the efficiency of the CMOS transducer. The plot on the right ofFIG. 2represents how a transducer (e.g., photodiode302ofFIG. 3) efficiently converts light energy (e.g., photon energy) into electrical energy.

In order to integrate the biochemical part of the assay with fluorescent detector104(FIG. 1B), capturing probes106(FIGS. 1C and 1D) are optically coupled to the CMOS detector using a fiber-optical faceplate (FOF) placed on top of CMOS integrated circuit101(FIG. 1A) as shown inFIG. 3in accordance with an embodiment of the present invention. Referring toFIG. 3, in conjunction withFIGS. 1A-D, in one embodiment, CMOS sensor array chip101has a thickness of approximately 530 micrometers which includes a silicon substrate301containing a photodiode302and readout circuitry303forming a pixel304. On top of silicon substrate301may include a dielectric layer305, such as silicon dioxide, containing metal “curtains”306A-B. Curtains306A-B may collectively or individually be referred to as curtains306or curtain306, respectively. Curtains306may be employed for further optical shielding against optical crosstalk by taking advantages of integrated circuit structures in addition to physical distance between neighboring pixels. In one embodiment, curtains306are composed of vias and metal layers (shown inFIG. 8A) and may encompass the entire photodiode302. Chip101may include any number of curtains306and the number of curtains306shown inFIG. 3is illustrative.

FIG. 3further illustrates an optical filter307(long-pass filter discussed above) residing on the top of chip101. In one embodiment, optical filter307has a thickness of approximately 2.1 micrometers. On top of long pass filter307may reside a fiber-optical faceplate308. In one embodiment, optical filter307includes layers of materials with a dissimilar refractive index.

In one embodiment, optical filter307is fabricated on the bottom of fiber-optical faceplate308and on the top of chip101. Optical filter307prohibits light scattering and guides the two-dimensional fluorescence signals along the vertical direction of its fibers. In one embodiment, the thickness of fiber-optical faceplate308may be between 0.5 millimeters and 3 millimeters which thermally isolates the 40-60° C. microarray assay from CMOS chip101and also creates adequate distance between the solution and chip101without any significant signal loss. The exposed surface of fiber-optical faceplate308may be polished glass (SiO2)309which is ideal for DNA capturing probe attachments using standard aldehyde-modified surfaces.

As stated above, on top of chip101resides a thin film dielectric optical filter307which blocks the excitation light, while only taking emission light from fluorophores107(e.g., Cy3 in this case) (indicated by arrows in silicon dioxide region305). As also stated above, on top of filter307resides fiber-optic faceplate308, which brings the bottom surface image (in this biosensor case, transducers integrated within CMOS chip101) to the top surface where biological analytes will be spotted in this case. Fiber-optic faceplate308may provide a good surface platform for DNA capturing probe attachments106while minimizing loss of signal due to the distance between the detectors (e.g., chip101on the very bottom) and the light generated by fluorophores107in the biological analytes. The biological analytes can be spotted on the top surface of fiber-optic faceplate308for the detection. As illustrated inFIG. 3, bio detection can be readily done on the single platform and the results are in digital numbers for further signal processing.

FIG. 3further illustrates capturing layer106used for capturing fluorescent labels that were excited by a fluorescent excitation signal310with a wavelength of approximately 532 nanometers.

CMOS image sensors may use a process referred to as “direct integration” to measure the emitted light from fluorophores107. In direct integration, the photocurrent generated in photodiode302is directly integrated (accumulated) on the photodiode capacitor. It is widely known in the art that direct integration can be carried out in an array format, where individual array components (i.e., pixels) measure light independently.

In an alternative embodiment, a capacitive transimpedance amplifier (CTIA) is used to create the photocurrent integrator as illustrated inFIG. 4.FIG. 4illustrates an embodiment of the present invention of pixel304. Some of the components of pixel304are shown in block diagram form in order not to obscure the present invention in unnecessary detail. A person of ordinary skill in the art would understand the workings of these components.

Referring toFIG. 4, in conjunction withFIG. 3, pixel304includes a capacitive transimpedance amplifier (CTIA)401. CTIA401includes a photodiode302, an amplifier403receiving an input voltage VR1and a capacitor404in the feedback mechanism of amplifier403. Further, CTIA401includes a switch405.

Unlike a direct integrator, the linearity of CTIA401is not limited by the photodiode junction capacitance voltage-dependency, and in addition, has a diode-independent well capacity set by feedback capacitor404. In one embodiment, feedback capacitor404is a 780 femtofarad poly-to-poly capacitor. Motivated by the advantages of CMOS digital-pixel-sensor (DPS) image sensors, an analog-to-digital converter406(ADC) (e.g., 14 bit analog-to-digital converter) has been integrated within the pixels. ADC406includes a comparator407coupled to a counter408where the output of counter408is a 14-bit digital output. Comparator407and counter408are controlled by a control logic409whose actions are coordinated by clock CLK2. Control logic409may receive control signals410used to program the actions of control logic409.

ADC406compares the output of CTIA401with the external reference voltage, VR2, to measure the time that the ramp reaches VR2using comparator407. To suppress the offset of comparator407, a chopper stabilized preamplifier411may be implemented with an overall voltage gain of 60 dB gain. Chopper stabilized preamplifier411includes mixers412,413as well as amplifiers414,415. Mixers412,413may also be referred to as “choppers” or “modulators.” Mixer412is the first chopper which modulates a signal to a higher frequency; whereas, mixer413is the second chopper which demodulates the signal back to baseband with an offset to cancel the low noise.

FIG. 4illustrates photodiode302converting emitted light into electrical signal (e.g., current). Then, CTIA401converts this current into voltage. The feedback switch405in CTIA401shown inFIG. 4provides extra large resistance. In a very low current input case, there will be limited finite CMOS switch resistance that prohibits the output of CTIA401going from the desired voltage. Switch405provides effectively much greater resistance than a single switch can provide while reducing the charge injections from switch405.

Chopper stabilized pre-amplifier411is used to amplify the output of CTIA401while reducing 1/f noise of the 1ststage of the amplifier used in pre-amplifier403of CTIA401. A non-overlapping CLK generator416is used to generate a clock signal to coordinate the multiplication of signal voltages by mixers412,413. The actions of CLK generator416are coordinated by clock CLK1.

The time that it takes for the voltage output of CTIA401to reach a certain voltage provides information about the amount of light detected by photodiode302. In turn, the amount of light detected by photodiode302is proportional to the abundance of the target analytes on the surface of fiber-optical faceplate308(FIG. 3). In one embodiment, all entire systems aforementioned are integrated in a single pixel.

FIG. 5is a graph500illustrating the measured sensitivity of the integrated microarray and the background current (Idc) present in the system as a function of temperature in accordance with an embodiment of the present invention. Referring toFIG. 5, graph500illustrates the measured sensitivity of the integrated microarray (as current in Amps)501as a function of the incident photon flux502(λ=532 nm), incident signal power503(λ=532 nm), and the estimated surface density of the fluorescent labels504(Cy3 label with a 3 mW excitation source). The measured dark current of each pixel304(FIG. 3) is approximately 12 fA at room temperature, which is much lower than the measured signals and therefore sufficiently low for typical fluorescence spectroscopy applications.FIG. 5illustrates the sensitivity of a transducer (e.g., photodiode302ofFIGS. 3 and 4) with CTIA401(FIG. 4).FIG. 5further illustrates the “dark current” (indicated as “Idc” inFIG. 5) that was measured at different temperatures (e.g., 20° C., 30° C., 40° C.). The dark current is the unwanted current that is typically generated by thermal agitation and leakage in photosensitive devices, such as photodiodes and charge-coupled devices. The integration time (indicated as “Tint”) indicates the time that was set in pixel304ofFIG. 4to measure the currents ofFIG. 5.

In addition to conventional microarray applications, system100(FIGS. 1A-D) is capable of real-time detection of fluorescent signals emitted from biological samples in solution.FIG. 6is a graph600illustrating the measured real-time DNA hybridization kinetics for three exemplary DNA target concentrations601,602,603in solution in accordance with an embodiment of the present invention.

Referring toFIG. 6, the target strand 5′-AGCAACATTTTGCTGCCG-3′ is labeled with Cy3 and the probe strand 3′-TCGTTGTAAAACGACGGC-5′ is labeled with a black-hole quencher. The sequences are complementary so that binding occurs.

DNA target concentration601corresponds to a 0.5 nmol target with a kH(normalized inverse of time constant) equal to 0.43×10−3sec−1. kHrefers to the normalized inverse of time constant which shows how much bindings occur in a given time. For example, the 0.5 nmol target case has more bindings in comparison to the 0.125 nmol target case in a given time. DNA target concentration602corresponds to a 0.25 nmol target with a kHequal to 0.31×10−3sec−1. DNA target concentration603corresponds to a 0.125 nmol target with a kHequal to 0.22×10−3sec−1.

The results shown inFIG. 6confirm that the measured rate of capturing is directly proportional to the analyte concentration. In one embodiment, the biosensor can detect hybridization kinetics in real-time. In a conventional microarray, washed and dried analytes labeled with fluorophores are typically used. Therefore, the researcher should wait until hybridization of DNA is done. Thus, only a steady-state response can be detected. However, in system100(FIGS. 1A-D), the solution can be spotted on the top of the sensor so that the detection can be performed during hybridization. Therefore, this integrated biosensor is able to detect real-time hybridization kinetic behavior. It is noted that this real-time detection does not require biosensors to be washed and waited until biological analytes get saturated. Detecting hybridization kinetic also provides information about the abundance of target analytes. In the illustrative case, all the capturing probes are labeled by fluorophores (e.g., Cy3) and the targets are labeled by quenchers, which extinct the light generated by fluorophores. The normalized results are shown inFIG. 6. The largest amount of target analytes is 0.5 nmol with the smallest time constant. FIG.6illustrates that larger target analytes in the solution represent the faster rising response. In summary, this type of integrated biosensor can be used for not only conventional detection (steady-state), but also real-time hybridization kinetics (transient).

The integrated biosensor of system100(FIGS. 1A-D) includes a fully integrated fluorescent-based microarray system. The achieved performances of this system in terms of sensitivity, compactness, versatility, and cost, satisfy the requirements of many biotechnology applications beyond DNA microarrays.

Though fluorescent based detection methods have been popularly used with microarrays, label-free detection of analytes using their intrinsic properties (e.g., charge, mass, absorption spectra) has generated a lot of interest in the research community. Label-free detection offers several advantages such as reduction in cost, omission of the molecular labeling process, feasibility of real time detection and ease of integration with standard CMOS processes.

Among the various techniques for label-free detection, impedance spectroscopy-based detection is the most compatible with current silicon-based very-large-scale integrated (VLSI) systems and integrated electronics. The concept behind this method is illustrated inFIG. 7in accordance with an embodiment of the present invention. In this method, an electrode surface701is first functionalized by immobilizing probes702(indicated as Y's attached to top of electrode surface701) on the electrode surface701. Electrodes701are placed in an aqueous solution703containing analytes704(analytes704are depicted inFIG. 7as various symbols, such as circles and triangles). Any number of analytes704may be present in aqueous solution703. When analytes704bind to probes702, the physiochemical characteristics of the electrode-electrolyte interface changes which subsequently results in changes in the impedance of the interface. It is known in the art that such changes in the impedance of the interface can be measured using an electronic sensor and the results can be related to the binding events, their frequency, and their abundance. As illustrated inFIG. 7, there are three important terms contributing to the impedance: Rb, the resistance of the solution and the impedance of the double layer represented by Rdand Cd. On the right ofFIG. 7, the Rb′, Rd′ and Cd′ indicate the modified values due to analyte binding. The spectroscopy-based detection is known as “impedance” spectroscopy-based detection to signify the interest in the changes in both the resistance and the capacitance of the interface as a function of excitation frequency. As impedance is a purely electrical quantity, one of the key advantages of this method is that all the signals are completely in the electronic domain. This attractive feature enables straightforward integration with standard CMOS processes.

In an illustrative design using the principles of the present invention, a 10×10 array of impedance sensors is integrated with the detection circuitry (shown inFIG. 10) for sensing the impedance of the interface between electrode701and electrolyte. Some innovative aspects of the design include the use of on-chip sensing electrodes and simple detection circuitry which is present underneath each of electrodes701, permitting large-scale integration and high packing density. Another innovative aspect is that the whole chip, including the sensors, may be fabricated using 0.35 μm standard CMOS process. Since CMOS process is widely used in digital integrated circuits, large scale production of “integrated impedance sensors” could lead to significant lowering of costs.

FIGS. 8A-Billustrate a pixel800with on-chip electrode701(FIG. 7) that is formed by creating an opening801in a passivation layer802covering the metal layers plus dielectric layer803in accordance with an embodiment of the present invention. Referring toFIGS. 8A-B, in conjunction withFIG. 7, in a CMOS process, active devices804are fabricated on a silicon substrate805, and 3-4 metal layers, which are used to route various signals, are stacked on top of these active devices804.FIG. 8Afurther illustrates vias806used to interconnect various layers. Passivation layer802covers the top layer of metal layers803in order to protect metal layers803and devices804underneath from the outside environment. An opening801in passivation layer802exposes the top layer of metal layers803to an aqueous solution of affinity-based biosensors, which is used as the sensing electrode701. Capturing probes702are attached to sensing electrode701. In one embodiment, passivation layer802has a thickness of approximately 2 micrometers.

FIG. 9illustrates a scanning electron microscope (SEM) photograph900of a surface of the fabricated pixel800(FIGS. 8A-B) in accordance with an embodiment of the present invention. Referring toFIG. 9, in conjunction withFIGS. 7-8, in one embodiment, electrode701is roughly of the size 40 μm×40 μm. In one embodiment, the edges of electrode701are rounded off in order to approximate a circle since a drop placed on top of electrode701tends to take a circular shape. In one embodiment, electrodes701are spaced approximately 50 μm apart. In one embodiment, the die area is approximately 2 mm×2 mm.

FIG. 10illustrates the excitation scheme of a chip1000in accordance with an embodiment of the present invention. Referring toFIG. 10, in conjunction withFIGS. 7 and 8, the surface of electrodes701is functionalized with capture probes702. Solution703containing analytes704is dropped on top of a detection circuitry1001residing on silicon substrate805of chip1000. In measuring the impedance of the system, an excitation source1002, which is a sinusoidal function generator, is used to generate sine waves in the frequency range from a 0.01 Hz to 500 MHz. In one embodiment, excitation source1002typically generates sine waves in the frequency range between 10 Hz to 10 MHz. The excitation signal may be applied to an Ag/AgCl electrode1003which is dipped in solution703. The impedance of the interface may be measured by detection circuitry1001.

FIG. 11illustrates the basic concept used in measuring impedance by detection circuitry1001(FIG. 10) in accordance with an embodiment of the present invention. Referring toFIG. 11, the input sinusoid Vin=A cos ωt1101is applied as the excitation source. The impedance Z1102represents the sum of the solution resistance and the interface impedance. The current flowing through Z is given by

This current is multiplied by two sinusoidal signals B cos ωt1103and B sin ωt1104, which have the same frequency of the excitation source. The path in which the multiplication with B cos ωt is called I path1105. The other path is generally referred to as the Q path1106. After multiplication by multipliers1107,1108and subsequent low pass filtering by low pass filters1109,1110, a signal proportional to cos θ is generated in I path1105, while the signal proportional to sin θ is generated in Q path1106. Using these two signals, it is possible to calculate both the magnitude and phase of the complex impedance denoted by Z inFIG. 11.

FIG. 12illustrates a circuit1200for the each individual pixel800(FIGS. 8A-B) in accordance with an embodiment of the present invention. Referring toFIG. 12, in conjunction withFIGS. 7,8and11, circuit1200includes a common gate input stage1201which includes a p-type transistor1202coupled to n-type transistors1203,1204. The gate of n-type transistor1203is coupled to a differential amplifier1205. Electrode701is connected to input node1206of common gate amplifier1205. Electrode701is coupled to the input voltage (Vin). The gate of n-type transistor1204is biased with a signal labeled “vbias_tail.” The negative input of common gate amplifier1205is coupled to node1206; whereas, the positive input of common gate amplifier1205receives an input voltage labeled “vin_cm.” The source of p-type transistor1202is coupled to power supply1207; whereas, the source of n-type transistor1204and Vinare coupled to ground1208.

Circuit1200additionally includes a mixer1209. Front-end amplifier1201and mixer1209may be implemented on-chip while other components (e.g., low pass filter and signal processing blocks) may be implemented off-chip in order to reduce the area of pixel800.

In one embodiment, common gate input stage1201presents a low input impedance below 100 ohms for the entire frequency range from DC to 50 MHz. This is achieved by further reducing the transconductance of input transistor1202using a simple differential amplifier1205as the gain-boosting circuitry. Another important function of differential amplifier1205is that it helps to set the DC potential at electrodes701. This helps in maintaining zero DC potential between the working electrode common to all the electrodes and the on-chip electrode. The current flowing through input stage1201is transferred using current mirrors1210to a set of double-balanced Gilbert-cell mixers1211. Double-balanced Gilbert cell mixers1211are adopted to suppress the component at the signal frequency caused by I and Q square waves1105,1106applied to mixer1209. To minimize mismatch, an exact replica of input circuit1201(circuit1212) is used carrying the same amount of current but with no input being applied. The only exception is that differential amplifier1213used in replica bias circuit1212has lesser current. Furthermore, resistors1214,1215are used in the load of mixer1209in order to minimize the flicker noise (1/f noise) at output1216.

Mixer1209additionally includes a current source identified by “Inix” that is inputted to current mirrors1210. The gate of the transistors1217,1218of current mirrors1210receives an opposite input voltage, vin+ and vin−, respectively. The drains of transistors1217,1218are coupled to the sources of transistors1219,1220,1221,1222of mixers1211. The gates of transistors1219,1222receive a positive oscillator voltage (identified as “VLO+”); whereas, the gates of transistors1220,1221receive a negative oscillator voltage (identified as “VLO−”). The negative and positive output voltage of mixers, identified as “Vout−” and “Vout+,” respectively, are coupled to resistors1214,1215, respectively. Resistors1214,1215are coupled to ground1208.

As discussed above, circuit1212is a replica of input circuit1201. Circuit1212includes a p-type transistor1223coupled to power supply1207. The gate of transistor1223is coupled to the gate of transistor1217. The drain of transistor1223is coupled to the source of n-type transistor1224. The gate of transistor1224is coupled to differential amplifier1213with its negative input coupled to the source of transistor1224. The positive input of differential amplifier1213receives the input voltage labeled “vin_cm.” The source of transistor1224is coupled to a current source labeled “lcg” which is coupled to ground1208.

The pixel level performance metrics using the components of circuit1200are shown in Table I.

TABLE IPerformance metrics of pixel 304Gain of the cell90 dBBandwidth10 Hz-50 MHzNoise at the output referred to the<0.4 nA rms over 100 Hz bandwidthinput currentInput impedance<100 ohm from dc to 50 MHz at allcornersMax input current20 μACurrent consumption210 μA at 3.3 V supply

In one embodiment, circuit1200has a current to voltage gain of 90 dB and the input referred noise current is less than 0.2 nA rms for 100 Hz bandwidth. In one embodiment, each pixel800consumes 210 μA of current with a 3.3V power supply. In one embodiment, the maximum input current for the circuit remains linear at 20 μA.

Although the systems are described in connection with several embodiments, it is not intended to be limited to the specific forms set forth herein, but on the contrary, it is intended to cover such alternatives, modifications and equivalents, as can be reasonably included within the spirit and scope of the invention as defined by the appended claims.