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US20100300899A1 - Active CMOS Sensor Array For Electrochemical Biomolecular Detection - Google Patents

Active CMOS Sensor Array For Electrochemical Biomolecular Detection Download PDF

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Publication number
US20100300899A1
US20100300899A1 US12/819,720 US81972010A US2010300899A1 US 20100300899 A1 US20100300899 A1 US 20100300899A1 US 81972010 A US81972010 A US 81972010A US 2010300899 A1 US2010300899 A1 US 2010300899A1
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electrochemical
chip
cmos
biomolecular
working electrodes
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Peter M. Levine
Kenneth L. Shepard
Ping Gong
Levicky Rastislav
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Columbia University in the City of New York
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Columbia University in the City of New York
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Assigned to THE TRUSTEES OF COLUMBIA UNIVERSITY IN THE CITY OF NEW YORK reassignment THE TRUSTEES OF COLUMBIA UNIVERSITY IN THE CITY OF NEW YORK ASSIGNMENT OF ASSIGNORS INTEREST (SEE DOCUMENT FOR DETAILS). Assignors: LEVINE, PETER M., GONG, PING, RASTISLAV, LEVICKY, SHEPARD, KENNETH L.
Publication of US20100300899A1 publication Critical patent/US20100300899A1/en
Assigned to NATIONAL INSTITUTES OF HEALTH (NIH), U.S. DEPT. OF HEALTH AND HUMAN SERVICES (DHHS), U.S. GOVERNMENT reassignment NATIONAL INSTITUTES OF HEALTH (NIH), U.S. DEPT. OF HEALTH AND HUMAN SERVICES (DHHS), U.S. GOVERNMENT CONFIRMATORY LICENSE (SEE DOCUMENT FOR DETAILS). Assignors: COLUMBIA UNIV NEW YORK MORNINGSIDE
Priority to US15/241,486 priority patent/US10718732B2/en
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    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N27/00—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means
    • G01N27/26—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means by investigating electrochemical variables; by using electrolysis or electrophoresis
    • G01N27/28—Electrolytic cell components
    • G01N27/30—Electrodes, e.g. test electrodes; Half-cells
    • G01N27/327—Biochemical electrodes, e.g. electrical or mechanical details for in vitro measurements
    • G01N27/3275—Sensing specific biomolecules, e.g. nucleic acid strands, based on an electrode surface reaction
    • G01N27/3277—Sensing specific biomolecules, e.g. nucleic acid strands, based on an electrode surface reaction being a redox reaction, e.g. detection by cyclic voltammetry
    • C—CHEMISTRY; METALLURGY
    • C12—BIOCHEMISTRY; BEER; SPIRITS; WINE; VINEGAR; MICROBIOLOGY; ENZYMOLOGY; MUTATION OR GENETIC ENGINEERING
    • C12Q—MEASURING OR TESTING PROCESSES INVOLVING ENZYMES, NUCLEIC ACIDS OR MICROORGANISMS; COMPOSITIONS OR TEST PAPERS THEREFOR; PROCESSES OF PREPARING SUCH COMPOSITIONS; CONDITION-RESPONSIVE CONTROL IN MICROBIOLOGICAL OR ENZYMOLOGICAL PROCESSES
    • C12Q1/00—Measuring or testing processes involving enzymes, nucleic acids or microorganisms; Compositions therefor; Processes of preparing such compositions
    • C12Q1/68—Measuring or testing processes involving enzymes, nucleic acids or microorganisms; Compositions therefor; Processes of preparing such compositions involving nucleic acids
    • C12Q1/6813—Hybridisation assays
    • C12Q1/6816—Hybridisation assays characterised by the detection means
    • C12Q1/6825—Nucleic acid detection involving sensors
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N27/00—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means
    • G01N27/26—Investigating or analysing materials by the use of electric, electrochemical, or magnetic means by investigating electrochemical variables; by using electrolysis or electrophoresis
    • G01N27/416—Systems
    • G01N27/48—Systems using polarography, i.e. measuring changes in current under a slowly-varying voltage
    • H10W20/42—
    • H10W20/435—
    • H10W20/4446—

Definitions

  • DNA sensing has broad application, for example, in genotyping, gene-expression studies, mutation detection, pharmacogenomics, forensics, and related fields in which genetic contents provide insight into biological function or identity.
  • Multiplexed DNA analysis can be performed in a laboratory environment using a “microarray”, a passive substrate (such as a glass slide) on which thousands of single-stranded DNA (ssDNA) “probe” molecules arranged in a regular pattern bind to (or “hybridize” with) fluorophore-labeled “target” molecules in an analyte solution.
  • ssDNA single-stranded DNA
  • Probes can be synthesized externally and then immobilized on the microarray through mechanical contact spotting or non-contact ink-jet printing, or can be constructed in situ using photolithographic techniques and solid-phase chemical synthesis. Hybridization occurs, for example, when the probe and target sequences are complementary to one another.
  • Microarray scanners employing laser sources that excite the fluorophores and photomultiplier tubes or CCD cameras that detect the emitted light, can measure surface-bound target densities down to 10 6 cm ⁇ 2 . Relative expression levels of bound targets at different array sites can then be quantified from the resulting image.
  • fluorescent techniques typically require labeled targets and bulky instrumentation, making them ill-suited for point-of-care applications.
  • Electrochemical sensing approaches to DNA detection rely on detecting changes occurring with hybridization at the interface between a metal “working” electrode (“WE”), functionalized with probe molecules, and a conductive target analyte solution.
  • WE metal “working” electrode
  • One example feedback circuit known as a “potentiostat” can be used to apply a desired potential across the WE interface and measure the resulting current.
  • probe-target binding can be detected, for example, by measuring changes in the direct (Faradaic) current flowing across the interface.
  • label-free sensing can be performed, for example, by measuring changes in displacement (non-Faradaic) current at the interface that occur due to surface-charge fluctuations.
  • CMOS complementary metal-oxide-semiconductor
  • FIGS. 24 a - c depict several sensor chip architectures.
  • FIG. 24 a depicts a sensor chip interfaced with one off-chip working electrode.
  • FIG. 24 b depicts an array of on-chip sensing elements connected individually to off-chip working electrodes.
  • FIG. 24 c depicts the integration of arrays of sensing elements with on-chip working electrodes, where all electrochemical reactions are carried out on the sensor chip surface.
  • the present subject matter includes devices and techniques for quantitative, real-time detection of DNA hybridization using active CMOS-integrated electrochemical microarrays.
  • the disclosed subject matter includes an array of generalized high-performance feedback control devices.
  • the array can include one or more working electrodes, local controllers, signal converters, counter electrodes, and control amplifiers.
  • the array can also include one or more reference electrodes, be controlled by a global controller and further connected to a clock generator and a digital counter.
  • a technique includes integrating one or more working electrodes, local controllers, signal converters, counter electrodes, and control amplifiers.
  • the formed arrays can be cleaned and packaged with protective packaging.
  • An electrochemical measurement technique for example a cyclic voltammetry technique, can be employed, and can include stimulating an electrochemical cell formed from a substance occurring on the surface of the array, and measuring the reactions, or stimulating and measuring electrochemical reactions of a biological substance on the surface of the array.
  • the array can be implemented in a CMOS process.
  • the presently disclosed subject matter also provides devices for electrochemical sensing of biomolecules, including an integrated circuit.
  • the devices can include one or more working electrodes on the integrated circuit, the one or more working electrodes configured to receive one or more biomolecular probes, a desired potential maintained through one or more reference electrodes, the one or more working electrodes configured to form a portion of one or more corresponding potentiostats, and a digitizing circuit on the integrated circuit configured to measure a signal indicative of a biomolecule sensing operation in real time.
  • the integrated circuit can include a complementary metal-oxide-semiconductor (CMOS) chip having a top metal layer operably connected to one or more vias for routing electrical signals.
  • CMOS complementary metal-oxide-semiconductor
  • the chip can be fabricated in a 2.5-V, 5-metal, 0.25- ⁇ m CMOS process.
  • the one or more working electrodes can include square, gold electrodes.
  • the one or more working electrodes can be adhered to the top metal layer with an adhesion layer.
  • the adhesion layer can include titanium.
  • the digitizing circuit can include a dual-slope analog-to-digital converter circuit.
  • the one or more working electrodes can be in contact with an electrolyte solution, and the electrolyte solution can include one or more target molecules.
  • Other embodiments can include techniques for electrochemical sensing of biomolecules including providing one or more working electrodes on an integrated circuit, a desired potential maintained through one or more reference electrodes, the one or more working electrodes configured to bind one or more biomolecular probes, the one or more working electrodes configured to form a portion of one or more corresponding potentiostats, and providing a digitizing circuit on the integrated circuit configured to receive a signal resulting from an electrochemical measurement operation to measure one or more aspects of a biomolecular reaction in real time.
  • the techniques can further include binding one or more biomolecular probes at the one or more working electrodes.
  • the electrochemical measurement operation can include cyclic voltammetry, linear-sweep voltammetry, square-wave voltammetry, ac voltammetry, ac impedance, or electrochemical impedance spectroscopy techniques.
  • the techniques can further include analyzing the biomolecular reaction to quantify surface target coverages.
  • the biomolecular reaction can be indicative of quantitative and specific detection of biomolecules.
  • the biomolecular reaction can be indicative of DNA sensing.
  • Further embodiments include techniques for manufacturing a CMOS-based array for electrochemically measuring a biomolecular reaction.
  • the techniques can include defining one or more openings in a passivation layer of a CMOS chip to expose a top metal layer, depositing an adhesion layer at one of the openings, the adhesion layer in electrical communication with the top metal layer, depositing a metal layer on the adhesion layer to form a working electrode at least one of the openings, the working electrode in electrical communication with the adhesion layer, the working electrode configured to bind one or more biomolecular probes, the working electrode configured to form a portion of a potentiostat, and electrically connecting an on-chip digitizing circuit to the working electrode, the digitizing circuit configured to measure an electrical signal resulting from the biomolecular reaction.
  • Defining the one or more openings can include using a wet etch process to selectively remove the top metal layer.
  • the deposition of the adhesion layer can include an electron-beam deposition procedure.
  • the techniques can further include encapsulating bond wires in a chemically resistant epoxy to shield the wires from exposure to an electrolyte.
  • a layer of polydimethylsiloxane can be included between the chip and a top plate, the polydimethylsiloxane layer preventing leakage of the electrolyte. Further techniques can include forming a reservoir above the chip to hold the electrolyte.
  • Some embodiments include techniques for electrochemical sensing of biomolecules including means for binding one or more biomolecular probes at one or more working electrodes of an integrated circuit, the integrated circuit including one or more working electrodes, a desired potential maintained through one or more reference electrodes, and a digitizing circuit on the integrated circuit configured to measure a signal indicative of a biomolecule sensing operation in real time, the one or more working electrodes configured to form a portion of one or more corresponding potentiostats, and means for performing an electrochemical measurement operation to measure one or more aspects of a biomolecular reaction in real time.
  • FIG. 1 shows a potentiostat according to one embodiment of the disclosed subject matter.
  • FIG. 2 shows a potentiostat according to one embodiment of the disclosed subject matter.
  • FIG. 3 depicts an illustration in accordance with an exemplary embodiment of the disclosed subject matter.
  • FIG. 4 depicts an architecture of the CMOS sensor array according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 5 depicts a die photograph according to the embodiment of FIG. 4 .
  • FIG. 6 depicts a circuit model according to another embodiment of the disclosed subject matter.
  • FIG. 7 a depicts the architecture of a dual-slope ADC according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 7 b depicts input and output voltage signals according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 8 a depicts a cross section of CMOS die before post-processing according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 8 b depicts a result of a wet etch process according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 8 c depicts an electrode according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 9 depicts an output noise spectrum according to an exemplary embodiment of the disclosed subject matter.
  • FIGS. 10 a and 10 b depict test results for an ADC according to an exemplary embodiment of the disclosed subject matter.
  • FIGS. 11 a and 11 b depict measured cell currents from an electrochemical redox reaction according to some embodiments of the disclosed subject matter.
  • FIG. 12 depicts measured results according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 13 depicts results according to another embodiment of the disclosed subject matter.
  • FIGS. 14 a and 14 b depict results according to further embodiments of the disclosed subject matter.
  • FIG. 15 depicts results according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 16 depicts components according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 17 depicts output according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 18 depicts results according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 19 depicts results according to another embodiment of the disclosed subject matter.
  • FIG. 20 depicts results according to yet another embodiment of the disclosed subject matter.
  • FIG. 21 depicts a procedure according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 22 depicts results according to an exemplary embodiment of the disclosed subject matter.
  • FIG. 23 depicts results according to another embodiment of the disclosed subject matter.
  • FIG. 24 depicts various CMOS sensor implementations.
  • CMOS electrochemical sensor arrays for biomolecular detection can eliminate the need for the bulky and expensive optical equipment used in fluorescence-based microarrays. Such a reduction in size and complexity paves the way for the use of electrochemical microarrays in point-of-care applications.
  • the design of a four-by-four array implemented in a standard 0.25- ⁇ m CMOS process augmented by post-processing to fabricate integrated electrochemically-compatible and biologically-compatible electrodes is employed.
  • Integrated potentiostat electronics and ADCs stimulate and measure electrochemical reactions occurring at the chip surface. Results from demonstrations of cyclic voltammetry measurements of redox species, characterization of DNA probe coverages, and quantitative and specific detection of DNA probe-target hybridization illustrate some bio-diagnostic capabilities of the chip.
  • a potentiostat is, for example, a feedback control device used to apply a desired potential to an electrochemical cell and simultaneously measure the movement of charge through the cell that accompanies electrochemical reactions occurring at an electrode-electrolyte interface.
  • One example potentiostat 200 used in a typical electrochemical setup, shown conceptually in FIG. 1 includes three electrodes immersed in an electrolyte: a WE 202 at which the reaction of interest occurs and to which biomolecular probes can be attached, a “reference” electrode (RE) 204 to hold the electrolyte at a known potential, and a “counter” electrode (CE) 206 .
  • the voltage V between WE 202 and CE 206 is adjusted to establish a desired cell input voltage V in between the WE 202 and RE 204 .
  • Direct current (I) can flow through the external circuit as measured by the ammeter (A) 210 as soluble redox species in the electrolyte, for example, donate electrons to the WE 202 and accept electrons from the CE 206 (through a “Faradaic” process).
  • a charging (displacement) current can flow as ions segregate to the WE 202 and CE 206 to form space-charge regions (through a “non-Faradaic” process).
  • the high-impedance voltmeter 208 attached to the RE 204 ensures that very low current flows through this interface, helping to maintain it at equilibrium.
  • the exemplary potentiostat circuitry in FIG. 1 can be implemented in one embodiment using standard electronic components as shown in FIG. 2 .
  • the control amplifier implemented as an operational amplifier (op amp) or operational transconductance amplifier (OTA) 250 on the right establishes the control loop while the integrator 252 on the left converts the current flowing through the WE 254 to a voltage for digitization and readout.
  • the high input impedance of the control amplifier 250 ensures that a very small current flows through the RE 258 .
  • This circuit can be implemented in a CMOS process and forms the basis of the exemplary integrated sensor array 260 .
  • cyclic voltammetry a low-speed, large-signal technique, is used to detect redox species, determine surface DNA probe concentration, and measure DNA probe-target hybridization is used.
  • V in e.g., expressed relative to a standard RE potential.
  • V RE is set to the initial voltage V i , ramped up to the vertex potential V v at the scan rate ⁇ , and then ramped down at the same rate until the final voltage V f (usually equal to V i ) is reached, as shown in FIG. 3 (measurement of an electrochemical cell containing a redox species in solution using CV), while the current flowing through the WE i WE is measured simultaneously.
  • V RE is viewed as a function of V in so that the potentials at which reduction or oxidation occur (indicated by the forward and reverse current peaks, respectively) can be easily discerned.
  • EIS electrochemical impedance spectroscopy
  • ac impedance single frequency “ac impedance”
  • the impedance of the electrode-electrolyte interface is determined by applying a small ac voltage signal to the cell and measuring the displacement current produced as shown in FIG. 15 . If a range of input frequencies is used, a Nyquist plot of the impedance can be made. Fitting this data to a circuit model of the interface (as shown in the figure) provides information about the “double-layer” capacitance C dl in this region. This capacitance is known to change as a function of surface charge at the WE.
  • an active CMOS sensor array is employed for DNA sensing.
  • One embodiment of an architecture of the active CMOS sensor array is shown in FIG. 4 and the accompanying die photograph is shown in FIG. 5 .
  • the chip can be fabricated in any conventional process, including but not limited to a 2.5-V, 5-metal, 0.25- ⁇ m CMOS process, measuring 5 mm-by-3 mm. Alternatively, a bipolar, BiCMOS, or SOI process can be used.
  • Each array site 402 includes a square Au WE 404 and a dual-slope ADC 406 with digital control circuitry to digitize the current flowing through the electrode.
  • the WEs in the top row have a side length of 100 ⁇ m with subsequent rows having lengths of 90, 80, and 70 ⁇ m.
  • Each row of four WEs shares a 15- ⁇ m-by-2500- ⁇ m CE 408 driven by a control amplifier 410 .
  • One embodiment includes the small-signal circuit model of the electrode-electrolyte interfaces in an electrochemical cell 600 , shown in FIG. 6 .
  • resistors R s1 , and R s2 represent the solution resistance between the WE 602 and RE 604 , and CE 606 and RE 604 , respectively.
  • a commercial potentiostat e.g., a CHI 700C-series, available from CH Instruments, Austin, Tex., with a 125- ⁇ m-diameter Au WE (e.g., available from ESA Biosciences, Chelmsford, Mass., USA), platinum-wire (Pt-wire) CE (e.g., available from Sigma-Aldrich, St. Louis, Mo., USA), and a standard silver/silver-chloride/3-M sodium-chloride (Ag/AgCl/3-M NaCl) RE (e.g., available from Bioanalytical Systems.
  • a commercial potentiostat e.g., a CHI 700C-series, available from CH Instruments, Austin, Tex., with a 125- ⁇ m-diameter Au WE (e.g., available from ESA Biosciences, Chelmsford, Mass., USA), platinum-wire (Pt-wire) CE (e.g.,
  • R 11 ranges from approximately 275 ⁇ in 1-M potassium phosphate buffer (PPB, pH 7.4) to 10 k ⁇ in 10-mM PPB.
  • Capacitors C WE , and C CE model the interfacial “double-layer” capacitance at the WE and CE, respectively.
  • the measured value of C WE is on the order of 1 to 100 ⁇ F cm ⁇ 2 , depending on the electrolyte composition and modification of the WE surface.
  • Resistors R ct1 and R ct2 model the charge-transfer resistance at each surface, with the former having a typical measured value between 100 k ⁇ and 1 M ⁇ .
  • the use of a three-electrode potentiostat configuration makes the measurement of the WE interface independent of the RE and CE impedances. However, knowledge of the R ct1 , R ct2 and C CE values is used to test amplifier stability when designing the active sensor array.
  • the potentiostat circuits can provide good noise performance and stability across a wide range of operating conditions.
  • the control amplifier ( 410 of FIG. 4 ) can be implemented with a two-stage, single-ended-output op amp having a dc gain of 87 dB.
  • the amplifier when CV is carried out using low-frequency input signals, can include input CMOS devices having a width and length of 4 min and 1 ⁇ m, respectively, to reduce the effect of flicker (1/f) noise on the output.
  • these large input-stage transistors can improve matching.
  • the integrated ADCs of the current embodiment can digitize currents in the 100-pA to 250-nA range that flow bidirectionally through the WE.
  • ADC operation at sampling rates up to 10 kHz can be sufficient to accurately reconstruct the cell current.
  • an example dual-slope ADC architecture, shown in FIG. 7 a can be used for this purpose because its dynamic range and sampling frequency can be adjusted.
  • the ADC 700 includes of an integrating amplifier 702 with a fixed on-chip 5-pF linear capacitor 704 , and track-and-latch comparator 706 .
  • NMOS switches are accompanied by “dummy” transistors to absorb injected charge during switching and digital counter with control logic.
  • FIG. 7 b shows the control signals, clocks, and integrator and comparator outputs for a typical ADC conversion cycle. After the integrator is reset during t rst , current I WE is integrated onto capacitor C f for a fixed time interval t 1 . During this period, the comparator clock ⁇ dk , which operates at a higher rate than the ADC sampling frequency f s , is gated in order to reduce the effect of switching interference on the integrator output.
  • ⁇ dk is enabled and the integrator output voltage V int is measured by the comparator during the period t d .
  • the capacitor is discharged using the appropriate constant current source I ref+ or I ref ⁇ (for example, implemented using a PMOS and NMOS current mirror, respectively, biased with off-chip resistors) until V int crosses the comparator threshold (set to analog ground voltage V agnd ), signaling the end of the time period t 2 .
  • a counter, operated on ⁇ dk digitizes the time intervals and sets the nominal ADC resolution.
  • the value of I WE can then be calculated from (t 2 /t 1 )/I ref .
  • Auto-zeroing in which the integrator offset is sampled prior to digitizing the cell current, can be performed by setting the ⁇ az signal appropriately. This procedure implements offset correction in addition to mitigating the effect of 1/f noise on the output.
  • the integrator op amp 708 can have the same or similar architecture as the control amplifier, with similar input transistor sizes to reduce 1/f noise. To maintain closed-loop stability, op amp 708 is compensated with a 25-pF metal-insulator-metal (MIM) capacitor C c in series with a 150- ⁇ polysilicon resistor R c between the first and second stages.
  • MIM metal-insulator-metal
  • a minimum phase margin of 45° is obtained. This technique also takes into account component process, voltage, and temperature (PVT) variation as well as the effect of the input impedances of other WEs and ADCs in the array during parallel operation.
  • PVT component process, voltage, and temperature
  • extensive computational transient analyses are carried out to ensure stability when large-signal inputs are applied to the electrochemical cell, as required in CV.
  • the total bias current required by the integrator op amp 708 can be, for example, 4 mA. With this bias, a temperature increase of 3.5° C. is observed when the chip is performing electrochemical measurements with all ADCs running simultaneously. No deleterious effects occur during biomolecular detection experiments due to this temperature increase.
  • the track-and-latch comparator following the integrator can make correct decisions with a clock frequency up to, for example, at least 50 MHz.
  • Transistors M 3 and M 4 separate the cross-coupled switching transistors at the output from the drains of input transistors M 1 and M 2 to reduce kickback interference.
  • the latter transistors have a width and length of 200 ⁇ m and 1 ⁇ m, respectively, to reduce offset due to mismatch.
  • Integrated transconductance amplifiers can be included to test the operation of the ADCs.
  • Each amplifier contains an op amp driving the gate of a large NMOS or PMOS transistor (depending on the desired current direction) with its inverting input connected to an off-chip 10-M ⁇ resistor that is connected to either the supply voltage or ground.
  • the feedback loop ensures that the applied voltage at the non-inverting input is established across the resistor. Currents up to 150 nA can then he forced into the ADC using full-rail input voltages.
  • post-processing of the fabricated chip can be performed to create an array of electrodes on the surface of the chip.
  • the electrodes can be formed using Au, but other suitable, conducting materials can be used, including platinum, indium-tin-oxide, highly doped silicon, conducting carbon paste and conducting polymers, silver, silicon dioxide, graphite, etc.
  • Au can be relatively electrochemically inactive in the presence of strong electrolytes and is easily modified by self-assembly of well-ordered monolayers of thiol, sulfide, or disulfide compounds through Au-sulfur bonding. As a result, thiolated ssDNA probes can be strongly bound to Au surfaces.
  • FIGS. 8 a - c demonstrate CMOS post-processing procedures applicable to some embodiments for fabricating a surface electrode array.
  • FIG. 8 a shows a cross section of a CMOS die before post-processing showing the top two aluminum (Al) metal layers at one electrode site.
  • FIG. 8 b shows the result of a wet etch process to remove the top Al metal layer.
  • FIG. 8 c shows the final electrode resulting from Ti—Au thin-film deposition.
  • CMOS technologies use stacked metal layers connected vertically by tungsten (W) “vias” to route electrical signals and form transistor interconnects and can be passivated with SiO 2 and Si 3 N 4 at the chip surface.
  • W tungsten
  • openings in the chip passivation layers are defined at the desired electrode sites above the top-metal aluminum layer in a manner similar to that for making bond pads or probe pads, as shown in FIG. 8 a .
  • the openings can be square openings while in alternative embodiments, other shaped openings can be used as appropriate, such as circular, rectangular, hexagonal, etc. Any suitable technique can be used for forming the openings, such as inductively-coupled plasma (ICP) etching.
  • ICP inductively-coupled plasma
  • Al is an electrochemically active metal that can, in some instances, become corroded and in some applications can be replaced as necessary.
  • the WEs are post-processed by first selectively removing the exposed Al metal at the electrode sites using, for example, a wet-etch process as displayed in FIG. 8 b .
  • a wet-etch process As displayed in FIG. 8 b .
  • a wet-etch process is a phosphoric-acid wet etch technique.
  • Other suitable techniques for removing the Al metal include sodium hydroxide wet etching or ICP etching. This is followed by the electron-beam deposition of 20 nm of an adhesion layer, for example, Ti. Any adhesion layer at any appropriate thickness can be added, including Cr.
  • one or more additional layers can be added on top of the adhesion layer, such as an Au layer as already described.
  • a 300-nm layer of Au is added, followed by a lift-off process to form the final electrodes in FIG. 8 c .
  • any appropriate material and thickness of the material can be used to form the electrodes, such as 200 nm of Au.
  • These electrodes can connect directly to the tungsten (W) vias of the CMOS back end.
  • W tungsten
  • the foregoing techniques of electrode fabrication require fewer lithographic steps than the construction of a “stepped” electrode structure and do not require the implementation of a full CMOS back-end process. Such a procedure results in lower fabrication costs and complexity because it is more easily implemented in existing CMOS technologies.
  • the CEs are fabricated in a similar way as the WEs.
  • the post-processed chip can be set in a 272-pin, 27 mm-by-27 mm, ball-grid array (BGA) package with the die surface exposed.
  • BGA ball-grid array
  • the metal bond wires connecting the input-output (I/O) pads along the chip perimeter can be shielded from electrolyte exposure through encapsulation in a heat-cured, chemical resistant epoxy such as Hysol FP4450HF and FP4451TD (available from Henkel, Dusseldorf, Germany) or EPO-TEK GE116 and GE120 (available from Epoxy Technology, Billerica, Mass., USA).
  • the packaged chip can be fastened in a surface-mounted printed-circuit board (PCB) socket with a top-plate.
  • PCB printed-circuit board
  • a 1-mm thick polydimethylsiloxane (PDMS) sheet having a square opening in the center to expose the chip surface, can be introduced between the chip and top-plate to prevent electrolyte leakage onto the PCB.
  • a 10-mL glass reservoir attached to the top-plate using epoxy, can hold the liquid analyte in contact with the chip surface.
  • An external Ag/AgCl-3-M NaCl reference electrode can be held in the reservoir with a Teflon cap.
  • FIG. 9 shows the output noise spectrum of the control amplifier of some embodiments over a bandwidth from 10 Hz to 21 kHz when the device is operated in a unity-gain configuration.
  • the corner frequency is located around 10 kHz.
  • the measured output noise voltage is 21.2 ⁇ V rms over the 10-Hz to 21-kHz band when the effect of 60-Hz line interfering tones and other interfering tones are neglected.
  • characterization of the dual-slope ADC at each sensor site is carried out using the on-chip test circuits.
  • the ADCs are operated at an f s of 2.5 kHz with ⁇ clk set to 3.5 MHz. Integration time t 1 is set to 23 ⁇ s and discharge time t 2 is allowed a maximum value of 315 ⁇ s, providing a nominal resolution of 10 bits. The remaining time during each conversion cycle is required to reset C f and select the appropriate reference current source. The maximum I in before integrator saturation using these settings is about 110 nA.
  • Reference currents I ref+ and I ref ⁇ are set to 15 nA and 18 nA, respectively.
  • DNL and INL values for the ADCs can be ⁇ 0.25 LSB and +0.38 LSB, respectively, with an LSB current of approximately 240 pA.
  • FIGS. 10 a - b demonstrate measured results from example ac linearity testing of the dual-slope ADC: (a) shows dynamic range and (b) shows output spectrum with a full-scale input at 103 Hz (resolution bandwidth is 0.31 Hz).
  • FIG. 10 a displays the signal-to-noise ratio (SNR) and signal-to-noise-and-distortion ratio (SNDR) of the ADC as a function of input current level. The lower end of the DR curve is fitted due to the inability to provide a sufficiently small ac voltage signal to the on-chip transconductance amplifiers for ADC testing.
  • SNR signal-to-noise ratio
  • SNDR signal-to-noise-and-distortion ratio
  • the DR is limited at the high end by integrator saturation.
  • a maximum effective resolution (ENOB) of 9 bits is achieved and is limited by the linearity of the test circuits.
  • a DR of greater than 10 bits is achieved from computational circuit analyses of the dual-slope ADC alone.
  • FIG. 10 b shows the result of an 8192-point fast Fourier transform (FFT) of the measured ADC output when a full-scale input is applied.
  • FFT fast Fourier transform
  • the redox species potassium ferricyanide, K 3 [Fe(CN) 6 ], is an exemplary compound used by electrochemists to study interfacial properties due to its highly-reversible behavior. At the appropriate potential, ferricyanide ions are reduced to ferrocyanide ions in the reaction Fe(CN) 6 3 ⁇ +e ⁇ Fe(CN) 6 4 ⁇ .
  • the use of the active CMOS sensor array for electrochemical sensing, CV measurements of 2-mM potassium ferricyanide in 1-M PPB (pH 7.4) are carried out. In these illustrations, the potential between the WEs in the array and the RE is scanned from +0.75 V to ⁇ 0.5 V and back at various rates while the cell current is observed at one WE.
  • UV ultraviolet
  • OES ultraviolet/ozone cleaning device
  • deionized water e.g., 18.2 M ⁇ cm
  • FIG. 11 a shows the cell current at one of the 100- ⁇ m WEs when an input scan rate of 72 mV/s used.
  • a zero-phase, low-pass FIR filter is used to post-process the raw data in MATLAB.
  • this 80-mV difference in peak potentials is relatively close to the theoretical value of 59 mV for a fully-reversible, single-electron redox process 120.
  • the magnitude of the current falls after each peak due to mass-transport limitations of the redox species to the WE surface.
  • ⁇ is increased from 24 mV/s to 480 mV/s and the peak reduction current is measured. It has been shown that the peak current i p at a planar electrode for a reversible reaction under diffusive control exhibits the relationship i p ⁇ AD O 1/2 C O * ⁇ 1/2 .
  • A is the WE area
  • D O and C O are the diffusion coefficient and bulk concentration of the oxidized species, respectively. Therefore, i p (measured from the charging current background) increases linearly with ⁇ 1/2 , as is observed in FIG. 11 b.
  • FIG. 12 shows the results from this measurement.
  • Surface roughness of the electron-beam-deposited Au layer causes the actual electrode area to be larger than its geometric (drawn) area by a factor of about 1.5. This has been accounted for in the results.
  • the Au WEs are functionalized with a monolayer of ssDNA probes.
  • CV measurements are then carried out in the presence of the redox species hexaamineruthenium (III) chloride (RuHex 3+ ) to determine probe surface density.
  • the redox-active counterion RuHex 3+ associates with the surface-immobilized DNA, causing the thermodynamics of the redox processes to be altered as a result. It is known that as probe coverage increases, the reduction potential for the reaction RuHex 3+ +e ⁇ RuHex 2+ shifts toward more negative values due, in part, to changes in local charge concentration, dielectric constant, spatial distribution, and solvation. Once calibrated, these measurements can be used to determine the probe surface coverage.
  • the chip is cleaned as described previously and incubated for 30 min in a 1-M MgCl 2 solution containing a known concentration of thiolated 20-mer DNA probe.
  • MCP mercaptopropanol
  • CV at a scan rate of 4 V/s is then carried out in 7 mL of 10-mM Tris buffer (pH 7.4) with 1- ⁇ M RuHex 3+ .
  • FIG. 13 shows the results from two different CV demonstrations at one 90- ⁇ m WE for DNA probe coverages of 1 ⁇ 10 13 cm ⁇ 2 and 4 ⁇ 10 12 cm ⁇ 2 .
  • These probe densities are obtained by incubating the chip in different concentrations of DNA probe solution and are verified using a set of calibration measurements on a commercial potentiostat with 125- ⁇ m-diameter Au WE.
  • the overall shape of the CV curves is different than those obtained in the foregoing section because, in this case, RuHex 3+ is a surface-adsorbed species which is not subject to mass-transport limitations.
  • Fc N-(2-ferrocene-ethyl) maleimide redox labels
  • Fc redox labels are known to be chemically stable and are electrochemically reversible. They are one example alternative to the use of radioactive isotopes for nucleic-acid sequencing and sensing.
  • Fc redox labels have been used to study the thermodynamics of DNA probe binding and to perform label-based detection of RNA hybridization, among various other applications.
  • labeled targets When labeled targets are introduced into an environment containing surface-bound probes, the number of target molecules that hybridize can be determined, in some instances, by an equilibrium that depends on such factors as the relative probe and target concentrations, buffer ionic strength, and temperature.
  • the amount of bound target is measured from the charge transferred due to the Fc reaction, with one electron contributed by surface-bound target on the WE.
  • This value can be determined, for example, from CV measurements by integrating the area enclosed by the Fc redox current after subtraction of background charging contributions, and then dividing the result (in Coulombs) by the magnitude of the electronic charge (1.602176 ⁇ 10 ⁇ 19 C) and electrode area.
  • This technique differs from “intercalation”-based approaches to DNA detection because the latter approach cannot provide a measure of the absolute amount of probe-target hybridization on the WE surface.
  • alternative “sandwich”-based assays cannot provide a quantitative measure of target coverage in real time because the redox-active molecules are added after hybridization has occurred.
  • sequences of the 3′-end thiolated 20-mer DNA oligonucleotide probes used in the CMOS biosensor array experiments are as follows: P 1 5′-TTT TAA ATT CTG CAA GTG ATJ-3′ (from Homo sapiens retinoblastoma 1 mRNA) and P 2 5′-TTT TTT TCC TTC CTT TTT TTJ-3′; where J represents a thiol group.
  • the target sequences are as follows: T 1 5′-FcCAC TTG CAG AAT TTA AAA-3′ and T 2 5′FcAAA AAG GAA GGA AAA AAA-3′. Sequences P 1 and T 1 , and P 2 and T 2 , are pairwise complementary, respectively, do not exhibit self-complementarity, and will not cross-hybridize. These model sequences demonstrate the functionality of the active CMOS biosensor array platform and allow verification against results from off-chip electrochemical experiments using similar sequences.
  • FIG. 21 illustrates one embodiment for the synthetic protocol for preparation of the N-(2-ferrocene-ethyl) maleimide redox label 2308 starting from ferrocene acetonitrile 2302 .
  • any well known procedure can be employed. Chemicals used in this procedure are available from, for example, Aldrich, St. Louis, Mo., USA.
  • N-ferrocene ethylamine 2304 is made by adding 100 mL dry ether to 0.68 g lithium aluminum hydride (LiAlH 4 , 18 mmol) with argon protection, followed by mild stirring.
  • the LiAlH 4 /ether mixture is placed in an ice bath and 1 g (4.5 mmol) ferrocene acetonitrile 2302 solid is added slowly. The ice bath is then removed and the reaction is allowed to proceed for 2 h, after which the reaction mixture is slowly quenched with concentrated NaOH while on ice. The pH of the final aqueous phase is measured to confirm basicity and thus the completion of quenching. Next, MgSO 4 is added slowly to the mixture to remove water, and is subsequently removed by filtering. The remaining mixture is loaded on a silica gel column, washed with ether, followed by washing with ethyl acetate.
  • the column is then eluted with the mixture ethyl acetate (75%)/methanol (20%)/triethylamine (5%). The entire elute is collected and concentrated under vacuum overnight to obtain 0.63 g dark-brown viscous liquid, with 60% yield.
  • the final product is formed by adding 250 mg (0.75 mmol) of N-ferrocene ethyl maleamic acid 2306 to a solution of 2 mL acetic anhydride containing 15% sodium acetate. The resulting mixture is heated to 70° C. for 3 h with stirring. H 2 O and NaHCO 3 are added to the mixture to neutralize remaining acetic acid. The reaction mixture is then extracted with 10 mL ethyl acetate three times. The organic phases are collected and dried over MgSO 4 . Next, the solution is concentrated, applied onto a silica gel column, and the product is eluted with ethyl acetate (15%)/hexane (85%). The pure fractions are pooled and the solvent evaporated off to give 63 mg yellow solids with a melting point of 98-100° C.
  • the DNA target labeling procedure is as follows.
  • Target oligonucleotides, thiolated at the 5′ end by the manufacturer, are first deprotected with dithiothreitol (e.g., DTT, available from Aldrich, St. Louis, Mo., USA) to liberate the thiol group.
  • dithiothreitol e.g., DTT, available from Aldrich, St. Louis, Mo., USA
  • a size-exclusion column e.g., PD-10, available from Amersham Biosciences, Piscataway, N.J., USA
  • Fc Fc overnight in 150 mM potassium phosphate buffer (e.g., PPB, pH 8.0, available from Fisher, Pittsburgh, Pa., USA) at a nominal DNA concentration of 24 ⁇ M and a 50-fold excess of Fc.
  • PPB potassium phosphate buffer
  • oligonucleotide purification cartridge e.g., OPC, available from Applied Biosystems, Foster City, Calif., USA
  • reverse-phase HPLC purification on, for example, a Beckman Coulter system (available from Beckman Coulter, Fullerton, Calif., USA) including a Model 125 high-pressure gradient HPLC pump and a Model 168 multi-wavelength diode array detector and equipped with a C-18 reverse-phase analytical column.
  • a flow rate of 0.5 mL min ⁇ 1 and a linear gradient of 12%-100% methanol in a solution containing 8.6 mM triethylammonium and 100 mM hexafluoroisopropyl alcohol in water (pH 8.1) are used.
  • the desired product is collected and evaporated under vacuum until dry.
  • PCR polymerase chain reaction
  • Detection of DNA probe-target hybridization with the active CMOS sensor array is carried out by scanning from zero to +0.35 V and back at ⁇ of 60 V/s. Due to the relatively high scan rate required, the sampling rate of the dual-slope ADCs is increased to 10 kHz with ⁇ clk set to 3.5 MHz. The fixed integration and maximum discharge times are 15 ⁇ s and 63 ⁇ s, respectively.
  • the typical measured SNDR is 43.7 dB at a current a level of 38 nA rms (corresponding to a level of ⁇ 6 dBFS) and the maximum DNL and INL are +0.22 LSB and +0.15 LSB, respectively.
  • the chip surface is cleaned as described previously.
  • a layer of ssDNA probes at each WE is constructed by incubating the surface of the chip in 1 M MgCl 2 solution containing 500 nM probe for 30 min. This provides a probe surface density of approximately 8 ⁇ 10 12 cm ⁇ 2 , determined from a set of calibration measurements performed off chip.
  • the chip is incubated in a 1 mM MCP solution for 90 mM which forms a self-assembled monolayer on the Au WEs.
  • the CV demonstrations are run in 7 mL of 1 M PPB (pH 7.4), for example, made by combining appropriate amounts of K 2 HPO 4 and KH 2 PO 4 in water.
  • FIG. 17 shows output from the ADC measured at one of the 100 ⁇ m WEs in the array, functionalized with probe P 1 , approximately 50 min after 50 nM of target T 1 is introduced to the device.
  • the current peaks due to the Fc redox reactions are evident above the charging current level, indicating hybridization between P 1 and T 1 .
  • the forward and reverse charging currents differ slightly because of hysteresis.
  • the WE interfacial capacitance is measured to be approximately 7 ⁇ F cm ⁇ 2 , which is in the range for an MCP-modified Au electrode at the exemplified DNA coverage and buffer ionic strength.
  • the charging current, before DNA target is added to the buffer is shown along with the sensor response from hybridization, 50 min after target addition, at one 100 ⁇ m WE.
  • the on-chip CEs can be bypassed and an external discrete operational amplifier (e.g., AD8628, available from Analog Devices, Norwood, Mass., USA), that can operate up to 3 V can be used to drive the RE and an off-chip Pt-wire CE.
  • an external discrete operational amplifier e.g., AD8628, available from Analog Devices, Norwood, Mass., USA
  • the control amplifiers can be designed using thick-gate-oxide transistors, which can be operated up to 3.3 V in the present CMOS process.
  • the integrated platform enables real-time quantification of surface-hybridized targets in a multiplexed fashion, allowing large-scale optimization of parameters affecting hybridization in diagnostic assays including probe coverage, target concentration, probe and target sequence, buffer ionic strength, and temperature.
  • concentration of DNA target in solution and the magnitude of the sensor output signal are described below.
  • the dashed line shows the Langmuir fit to this equilibrium isotherm which yields a binding constant K a of approximately 1 ⁇ 10 8 M ⁇ 1 .
  • Error bars for the four lowest target concentrations indicate the standard deviation from three separate illustrations. Assuming the kinetics of hybridization between probe P and target T to form the DNA duplex D follow the reversible reaction
  • the equilibrium association constant K a can be determined using
  • K a is found to be approximately 1 ⁇ 10 8 M ⁇ 1 . This value of K a falls in the range determined in other work involving surface-based assays (10 7 -10 9 M ⁇ 1 ).
  • the hardware detection limit of the device is determined by the lowest measurable current.
  • the hardware sensitivity limit can be evaluated from electronic measurement of the dynamic range of the sensor ADCs (when operated at a 10 kHz sampling rate). Such measurements indicate that a current of 550 pA can be detected with a signal-to-noise ratio of three. This sets a detection limit of approximately 4 ⁇ 10 10 cm ⁇ 2 (or equivalently, a 50 pM target concentration given the calculated value of K a above) assuming the maximum redox current I max from the Fc reaction can be expressed using
  • I max n 2 ⁇ F 2 4 ⁇ R ⁇ ⁇ T - v ⁇ ⁇ A ⁇ ( S T N A ) , ( 3 )
  • n is the number of electrons transferred (one for the Fc reaction)
  • F is the Faraday constant
  • R is the molar gas constant
  • T is the absolute temperature
  • ⁇ is the CV scan rate.
  • alternative protocols for label-based detection coupled with the implementation of a higher resolution ADC, significantly improves the detection limit of the CMOS electrochemical sensor array, simultaneously.
  • use of multiple (e.g., electroactive dendrimer) labels can be used to boost signal per hybridized target.
  • employing background subtractive measurement techniques such as ac voltammetry or square-wave voltammetry can be used to reduce background currents and amplify the desired current from the redox activity of the labels.
  • Alternative protocols for label-based detection include the use of biotin-streptavidin chemistry or added intercalater molecules.
  • SPR surface plasmon resonance
  • QCM quartz crystal microbalance
  • cantilever sensors are capable of performing real-time DNA sensing.
  • electrochemical sensing techniques include simpler hardware and facile CMOS integration without surface micromachining or more complex post-processing as with cantilever or QCM fabrication.
  • FIG. 19 demonstrates the measurement of real-time kinetics in one embodiment where 60 nM of T 1 is hybridized to complementary P 1 .
  • This CV measurement is taken at one of the 100- ⁇ m WEs, with a scan repeated every 5 minutes and the cell potential held at 0 V between scans.
  • the measured data is fit to a first-order rate equation (dashed line) following Langmuir kinetics.
  • Inset shows the results from each CV scan over time. An increase in the area of the redox target peak is evident over time. The maximum extent of hybridization, reached after about 35 min, is about 6.8 ⁇ 10 12 cm ⁇ 2 .
  • multiplexed and specific detection are accomplished using the CMOS biosensor array by functionalizing the chip with two distinct probes and hybridizing each with its complementary target.
  • Probes P 1 and P 2 are spotted on four different WEs each using a fluid microinjection device (e.g., 1M-300, available from Narishige, East Meadow, N.Y., USA) capable of delivering nanoliter volumes of probe solution to the electrode surface.
  • a fluid microinjection device e.g., 1M-300, available from Narishige, East Meadow, N.Y., USA
  • FIG. 20 also shows the measured target coverages at sites A and B as a function of time elapsed since target addition.
  • the values of ⁇ for the hybridization processes are approximately 540 s and 740 s at site A and B, respectively.
  • the slight shift of the data relative to the origin is attributed to mass-transport limitations at the early times of hybridization.
  • the chip is cleaned as described previously and is then incubated in a 0.50- ⁇ M solution of 20-mer DNA probe, followed by incubation in MCP.
  • the sensor array is operated in 1-M PPB and the baseline current level at a 100- ⁇ m WE is first measured as shown in FIG. 14 a.
  • FIG. 14 b displays the resulting CV curve using the single-mismatch target sequence along with the curve obtained from the fully-complementary target for comparison. Because the target containing the SNP has less affinity for the probe, a smaller fraction of probe is hybridized, as can be seen from the reduced signal level in the figure. The density of hybridized probe and target in this case is 2.38 ⁇ 10 12 cm ⁇ 2 .
  • the density of the current CMOS biosensor array is approximately 250 cm ⁇ 2 . In other embodiments, the density can be increased to more than 6000 cm ⁇ 2 for the same 100 ⁇ m ⁇ 100 ⁇ m WE area by optimizing the physical layout of the on-chip electronics. This density would be comparable to existing, commercial detection devices, while additionally incorporating the full potentiostat sensing electronics on chip not present in the commercial devices.
  • the CMOS biosensor array can be used in clinical gene expression samples that have traditionally been analyzed with fluorescence-based arrays.

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