ARTICLE pubs.acs.org/ac
Portable Device Based on Chemiluminescence Lensless Imaging for Personalized Diagnostics through Multiplex Bioanalysis Aldo Roda,*,†,‡ Mara Mirasoli,† Luisa Stella Dolci,† Angela Buragina,† Francesca Bonvicini,§ Patrizia Simoni,‡,|| and Massimo Guardigli† †
Department of Pharmaceutical Sciences, University of Bologna, Via Belmeloro 6, 40126 Bologna, Italy National Institute for Biostructures and Biosystems (INBB), Viale Medaglie d'Oro 305, 00136 Rome, Italy § Department of Haematology and Oncological Sciences “L. e A. Seragnoli”, Microbiology Section, and Department of Clinical Medicine, University of Bologna, Via Massarenti 9, 40138 Bologna, Italy )
‡
ABSTRACT: A simple and versatile analytical device designed to perform, even simultaneously, different types of bioassays has been developed and optimized. A transparent microfluidics-based reaction chip, where analytes were quantitatively detected by means of biospecific reactions and chemiluminescence detection, was placed in contact with a thermoelectrically cooled CCD sensor through a fiber optic taper. Such a lensless contact imaging configuration combined adequate spatial resolution and high light collection efficiency within a small size portable device. The miniaturization of the reaction chamber ensured short analysis times (in the minutes range), while the use of chemiluminescence detection provided wide signal dynamic range and high detectability, down to attomole levels of protein and femtomole levels of nucleic acid analytes. A model hybrid panel test was realized by combining an enzyme assay for alkaline phosphatase activity, a nucleic acid hybridization assay for Parvovirus B19 DNA, and an immunoassay for horseradish peroxidase as a model antigen. The successful simultaneous quantification of the three targets demonstrated that a range of analytes, from enzymes to antigens, antibodies, and nucleic acids, can be measured in a single run, thus enabling the realization of a complete, personalized diagnostic panel test for early diagnosis of a given disease and patient follow-up.
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he advent of personalized medicine requires the development of “personalized diagnostic” approaches, i.e., the definition and the subsequent detection for each patient of the panel of the most powerful and specific biomarkers for the diagnosis of a given pathology and/or the follow-up of a given drug therapy.1,2 For such applications, highly flexible analytical devices should be designed, able to perform multiplexed analysis for determining a wide range of analytes, from disease biomarkers of a different nature (enzymes and enzyme substrates, antigens, antibodies, and nucleic acids) to blood levels of therapeutic agents. Continuous advances in microfluidics have led to the development of lab-on-chip devices, paving the way for miniaturized, self-standing analytical systems.3,4 Nevertheless, only recently technological solutions for combining different biospecific reactions (e.g., enzymatic, immunological, and nucleic acid hybridization reactions) in the same device are being devised.5,6 Chemiluminescence (CL) is considered the most suited optical detection technique for developing miniaturized, highly sensitive analytical devices.710 Even though the quantum efficiency of CL reactions is usually low (in the order of 0.01 or less), the use of enzyme labels such as horseradish peroxidase (HRP) or alkaline phosphatase (ALP) ensures high sensitivity thanks to the noteworthy signal amplification obtained in the presence of an excess of CL substrate.1012 Because of the wide dynamic range of the CL measurements (up to 6 orders of magnitude), the target analyte can be detected in a broad range of r 2011 American Chemical Society
concentrations, from femtomolar to millimolar levels, without the need of sample dilution. Finally, the absence of an excitation light source as in fluorescence measurements makes CL detection less sensitive to interferences due to sample components (the only background signal derives from the instrumental noise) and allows a flexible geometry of the measurement cell (e.g., wells, flow channels, microarrays of immobilized probes, etc.). Although several lab-on-chip devices relying on CL detection have been developed,1316 detection is usually performed by employing either laboratory instrumentation, thus not fulfilling the full portability of the device, or by miniaturized light sensors integrated in the chip, which are often not adequate for ultrasensitive detection of weak CL signals. Fluorescence or CL imaging detection employing charge-coupled devices (CCD) or complementary metal oxide sensors (CMOS) has been also proposed for low-cost portable devices, also providing the flexibility of microchannel designs.1719 Usually, lens-based optics are employed to achieve high spatial resolution. Nevertheless, these optical systems provide limited light collection efficiency (depending on the lens’s numerical aperture), and therefore, the lensless contact imaging approach, in which the object to be imaged (e.g., an array of microspots immobilized on Received: February 10, 2011 Accepted: March 14, 2011 Published: March 24, 2011 3178
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Figure 1. Scheme of the portable device for CL-based hybrid assays based on lensless imaging detection. Left: scheme of the reaction chip and of its different fluidic arrangements for enzyme activity assays in solution or binding assays. Center: scheme of the components of the biosensing device. Right: photo of the CCD camera assembled with the fiber optic taper.
a thin transparent support) is directly in contact with the imaging sensor, has been proposed.2022 In this work, we developed a new portable biosensing device for ultrasensitive and fast CL-based analysis, composed of a transparent microfluidics-based reaction chip that is imaged by a thermoelectrically cooled CCD sensor through a fiber optic taper, which coherently addressed the emitted photons to the sensor pixels in a compact geometry. Analyses were performed by exploiting biospecific interactions (e.g., enzymatic reactions, antigenantibody binding, nucleic acids hybridization): arrays of immobilized biospecific probes were employed to perform flow-assisted binding assays, while enzyme activities were measured in solution using suitably designed flow cells. The CCD camera was controlled by a portable PC, by means of which light emission intensity and 2D distribution data were easily acquired and processed. The analytical performance of the CL imaging system was evaluated to set up and optimize the bioassay formats. As a proofof-concept, the suitability of the device for diagnostics was demonstrated by performing models of the most common clinical chemistry analyses, i.e., enzyme activity assays (measurement of ALP activity in solution), immunoassays (measurement of HRP, employed as a model antigen, using an array of immobilized antiHRP antibodies), and nucleic acids hybridization assays (measurement of a target Parvovirus B19 nucleic acid sequence by an array of immobilized capture oligonucleotide probes). The proposed device allows short analysis time due to miniaturization and high flexibility of assay design and format, as well as a potential costs reduction. Its key advantage is that the different types of assays described above can be performed simultaneously exploiting microarray configurations and/or using different CL chemistries and labels. This multiplexing capability will permit the development of chips for the detection of panels of diagnostic biomarkers, e.g., for performing the diagnosis of an infectious disease by detecting both the pathogen nucleic acids and the host antibodies produced in response to the infection or to assess the liver function by combining routine clinical chemistry analyses (bilirubin, cholesterol) with the measurement of serum enzyme activities (alkaline phosphatase, aspartate, and alanine aminotransferase) and the detection of a present or past hepatitis viral infection.23
Thanks to the flexibility of chip design, chips to evaluate panels of biomarkers could be also specifically developed for a given patient to perform “personalized medicine” on the basis of genetic approaches.
’ EXPERIMENTAL METHODS Reagents. Horseradish peroxidase, anti-HRP antibody, bovine serum albumin (BSA), and 3-glycidoxypropyl trimethoxysilane (GOPS) were purchased from Sigma-Aldrich (St. Louis, MO). The SuperSignal ELISA Femto HRP CL substrate was obtained from Thermo Fisher Scientific, Inc. (Rockford, IL). The Roche Alkaline Phosphatase Calibrator for automated systems (460 IU/L ALP) was kindly provided by the Central Laboratory of the St. Orsola-Malpighi Hospital (Bologna, Italy). The LumiPhos Plus ALP dioxietane-based CL substrate was purchased from Lumigen (Southfield, MI). The HRP-labeled antidigoxigenin Fab fragments, dNTPs mix, Dig-dNTPs mix, FastStartTaq polymerase, and PCR ELISA (DIG-detection) buffers (hybridization buffer, denaturation buffer, conjugate dilution buffer, washing buffer) were purchased from Roche Diagnostics GmbH (Mannheim, Germany). All primers (forward primer 50 CGCCTGGAACACTGAAACCC-30 ; reverse primer 50 -GAAA CTGGTCTGCCAAAGGT-30 ) were obtained from Eurofins MWG Synthesis GmbH (Ebersberg, Germany). The Wizard SV Gel and Clean-Up System for purification of PCR products were from Promega Corporation (Indianapolis, IN). All other reagents were of analytical grade and were used as purchased. Biosensing Device. The biosensing device (Figure 1) consisted of a CCD camera modified for lensless CL imaging detection, a reaction chip, a housing for reaction chip positioning and light shielding, and external syringe pumps for fluids delivery. The reaction chip comprises a fluidic system containing the flow channels and connections required to perform the various analytical steps (e.g., flowing samples and reagents, washings, etc.). CCD Imaging Detector. The CCD imaging detector was built from a MZ-2PRO CCD camera (MagZero, Pordenone, Italy) equipped with a Sony ICX285 progressive scan monochrome CCD image sensor (1360 1024 pixels, pixel size 6.45 6.45 μm2) and a 16 bit analog-to-digital (A/D) converter. To minimize 3179
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Analytical Chemistry the instrumental thermal noise, the CCD sensor is thermoelectrically cooled by a double Peltier cell. The cooling system is driven by an external control box, which can be powered either by a 220 to 12 V ac/dc power adapter or a 12 V battery. The camera is computer-controlled via a USB 2.0 interface by a software that allows one to set the camera operation parameters and perform data acquisition. Images are recorded in the flexible image transport system (FITS) format, a digital file format commonly used in astronomy that is compatible with most scientific image analysis software. To perform lensless imaging, a round fiber optic taper (25/11 mm size, Edmund Optics, Barrington, NJ) was placed in contact with the CCD sensor. A homemade aluminum adapter with a fitted rubber O-ring was used to firmly hold the fiber optic taper in the right position. Finally, a chip holder with cover assured reproducible positioning of the chip during the measurement and provided shielding from ambient light. Reaction Chip. Parallel channels were engraved in a polycarbonate element (length 80.0 mm, width 25.0 mm, height 5.0 mm) with different geometries depending on the assay type (Figure 1). This element was coupled with a 170 μm-thick, 60.0 24.0 mm2 microscope borosilicate glass slide (Chemil s.r.l., Padova, Italy) using an adhesive foil, in which channels were also cut. For binding assays (immunoassays or gene hybridization assays), in which capture probes were immobilized on a functionalized glass slide along the channel pathway (see Immobilization of Biospecific Probes), the channel dimensions were 45.0 mm (length), 2.0 mm (width), and 0.5 mm (height), with a channel-to-channel separation of 6.0 mm and a channel volume of 60 μL. For enzyme activity assays in solution, cylindrical flow cells (diameter 3.0 mm, depth 3.5 mm, volume 25 μL) were created. Each cell was connected to a channel similar to those described above (20.0 mm length), while a second fluidic connection was at the top of the cell. Flow System. PEEK tubings were used to connect the channels in the reaction chip to KDS100 syringe pumps (KD Scientific, Holliston MA) for fluid handling. Flow rates were in the range 10400 μL/min, depending on the analytical protocol. Chemiluminescence Detectors. In addition to the CCD sensor described above, a CMOS sensor (Micron Imaging MT9M001, Micron Technology Inc., Boise, ID), equipped with a 1/2 in. monochrome CMOS sensor (1280 1024 pixels) with a quantum efficiency of 52% at 550 nm, and a Night OWL LB 981 research luminograph (Berthold Technologies GmbH & Co KG, Bad Wildbad Germany) with conventional lens-based optics and a highly sensitive, back-illuminated, double-Peltier-cooled CCD camera2426 were also used. Assessment of Light Detectability. To evaluate the performance of the lensless CCD imaging device, mini microtiter plates were obtained by cutting 4 3-well arrays from 384-well black polystyrene microtiter plates with a clear optical bottom (μClear 384 Well Small Volume LoBase Polystyrene Microplates, gently provided by Greiner Bio-One GmbH, Frickenhausen, Germany). These plates have conical round wells (bottom diameter ∼2.0 mm, maximum working volume 25 μL) sealed at their bottoms with a 75 μm optically clear film. Thus, the mini microtiter plates can be imaged either from the bottom with the lensless CCD device or from the top with the LB 981 luminograph. Calibrations curves for HRP in solution were produced by dispensing in triplicate in the wells 1 μL of HRP solutions with concentrations ranging from 0.9 to 60 pmol/L. Then, 20 μL of HRP CL substrate
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was added and the photon emission was immediately imaged using a 1 min acquisition time. Immobilization of Biospecific Probes. Microscope borosilicate glass slides were cleaned in piranha solution (concentrated H2SO430% H2O2, 3:1 v/v) for 30 min, thoroughly rinsed with ethanol 90%, ethanol 30%, and distilled water, and dried at 65 °C for 15 min. Then, the glass surface was treated with GOPS (10% v/v in toluene) for 2 h at room temperature, rinsed twice in toluene, sonicated in toluene for 10 min to remove noncovalently absorbed silane compound, and dried at 80 °C for 30 min. Spots of biospecific probes (antibodies or NH2-modified oligonucleotide probes) were deposited on the functionalized glass slide corresponding to the fluidic channels engraved in the polycarbonate element using a manual spotter (VP 470 Glass Slide Indexing System, Bio Gene, United Kingdom), which allows depositing spots of 2030 nL volume and 600800 μm diameter, depending on the solution viscosity and surface tension. Model Application 1: Measurement of Alkaline Phosphatase Activity in Solution. Alkaline phosphatase activity was determined by employing a 1,2-dioxetane phosphate-based CL substrate.27 Enzyme solutions in 150 mM NaCl with enzyme activities ranging from 10 to 230 IU/L of ALP were mixed 1:5 (v/v) with the CL substrate employing a tee junction and injected in the reaction chip. Then, the flow was stopped and the CL signal was imaged for 10 min using 1 min sequential acquisitions. Model Application 2: Immunoassay of HRP Antigen with Immobilized Anti-HRP Antibody Microarrays. Arrays of spots of anti-HRP antibody (0.1 μg/spot) were prepared by deposition on the activated glass slide of an antibody dilution in phosphate buffered saline (PBS). After incubation for 1 h at room temperature in a humid chamber, the slides were washed with PBS (3 5 min) and then treated to reduce nonspecific binding by soaking in blocking solution (3% BSA in PBS 0.1 M, pH 7.6) for 2 h under gentle agitation at room temperature. The functionalized glass slide was coupled with the polycarbonate fluidic element as described above. To obtain a calibration curve, 200 μL of HRP antigen solutions in PBS at different concentrations (from 1.5 103 to 570 pmol/L) were flowed (50 μL/min) through the channel at room temperature. The channel was filled with the HRP CL substrate after washing with PBS, and the CL signal was immediately imaged. Model Application 3: Determination of Parvovirus B19 DNA with Immobilized Oligonucleotide Probe Microarrays. The Parvovirus B19 target DNA sequence and the 50 -amino modified specific capture probe were produced by amplification from a standard plasmid target28 employing digoxigenin-labeled nucleotides (Dig-dNTPs mix) and 50 -NH2-C12-modified forward primers, respectively. Briefly, 10 ng of plasmid was added to a reaction mixture (50 μL final volume) containing 0.1 mM dNTPs mix, 0.1 μM of each primers, 2 units of FastStartTaq polymerase, and the buffer supplied by Roche with MgCl2 at the final concentration of 2 mM. After an initial denaturation step at 95 °C for 5 min, 40 cycles were performed under the following conditions: 95 °C for 30 s, 55 °C for 30 s, and 72 °C for 30 s, with a final extension step at 72 °C for 5 min. Amplified products were purified by using the Wizard SV Gel and Clean-Up System, and their amounts were estimated by measuring the absorbance at 280 nm. The capture probe was heat-denatured (5 min at 95 °C, followed by 5 min at 0 °C), and then probe spots were immobilized on the glass activated slide (1 pmol/spot). To block nonspecific absorption, the glass slide was soaked in blocking solution for 1 h under gentle shaking at room temperature. To obtain a calibration curve, 3180
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Analytical Chemistry 1 μL of digoxigenin-labeled Parvovirus B19 target DNA solutions at different concentrations (from 0.04 to 4 μmol/L) was denatured by adding 20 μL of denaturation buffer and, after addition of 180 μL of hybridization buffer, flowed (10 μL/min) in the channel. After washing (2 mL of washing buffer at 400 μL/ min), 200 μL of HRP-conjugated antidigoxigenin antibody diluted 1:12 000 (v/v) in conjugate dilution buffer was flowed at 20 μL/min. After washing with PBS, the channel was filled with the HRP CL substrate and the CL signal was immediately imaged. Chemiluminescence Image Analysis. Quantitative analysis of the CL images was performed using the WinLight software v. 1.2 (Berthold Technologies GmbH & Co. KG). For all measurements, the mean photon emission was calculated in the areas corresponding to the biospecific probe spots, wells, or flow cells. For production of calibration curves, the mean CL signals were corrected by subtracting the background signal measured in the absence of the target analyte.
’ RESULTS AND DISCUSSION Biosensing Device Design. The MZ2-PRO CCD camera is
particularly suited for implementation in a portable analytical device. With a design for outdoor application, it is lightweight (about 400 g) and can be battery operated. The CCD camera represents the most expensive component of the device: it was bought at a cost of about $2,000, but the price of similar devices will presumably decrease in the near future. The high quantum efficiency of the Sony ICX285 CCD sensor (higher than 50% in the 420680 nm interval) and the thermoelectrical cooling system (which allows the operation of the CCD sensor at low temperature, down to approximately 20 °C) permit the detection of very weak luminescent emissions. Moreover, no special care for thermal insulation, except sealing of the CCD sensor and electronics to prevent the fogging phenomena, was required. Therefore, the fiber optic taper could be easily implemented in the device without degrading the performance of the cooling system. The taper was necessary to perform lensless imaging without a direct contact between the reaction chip, in which the biological reactions must take place at room temperature, and the cooled CCD sensor. Furthermore, it allowed one to enlarge the active measurement area by a factor of 2.3 in comparison to the actual size of the CCD sensor (i.e., from 9.0 6.7 mm2 to 20.7 15.4 mm2), thus permitting use of larger reaction chips. Further improvements can be also envisaged in the development of this device, even though they were actually not implemented. For example, since the rate of an enzyme-catalyzed reaction depends on the temperature and the CL detection principle relies on enzymatic reactions, to obtain reliable and reproducible analytical results (especially when operating at a variable environmental temperature) temperature control elements should be integrated in the reaction chamber.29 Moreover syringe pumps could be replaced by more compact and less power demanding devices for fluid handling, such as miniaturized piezoelectric or diaphragm pumps. As concerned preanalytical sample treatment, the high sensitivity and selectivity obtained by combining biospecific reactions with CL detection will avoid the need of complex sample processing or preconcentration procedures. Immunoassays and enzyme activity assays could be performed directly on diluted blood samples after separation of blood cells, while for gene probe assays conventional or isothermal PCR will be used for amplifying the target sequences.
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Figure 2. Temperature dependence of the background signal of the CCD sensor; data (9) represent the mean value ( standard deviation (SD) of the background signal evaluated over the entire sensor area using a 1 min acquisition time. For reference, the signal measured for a spot in an inkjet printed mask (0), corrected for the background, is also reported.
Several devices for performing these operations have been already developed for microfluidics-based bioanalysis30 and could be implemented in our analytical system. Characterization of the CCD Device. The performance of the CCD device has been investigated by using simple CL systems (such as HRP standard solutions) or models simulating CL light sources. Temperature Effect. The background signal due to the dark current, caused by the temperature-dependent thermal generation of electrons, is one of the main factors that reduces the sensitivity of CCD sensors. Since the dark current usually shows a nearly exponential temperature dependence, CCD sensors are thermoelectrically or cryogenically cooled to reduce thermal noise and increase sensitivity. To confirm this behavior for the Sony ICX285 CCD sensor, we measured the background signal at different temperatures. As shown in Figure 2, the background of the CCD is drastically reduced in going from 25 °C to 10 °C. Most remarkably, its variability is also significantly reduced, being the standard deviation (SD) of the background signal measured at 10 °C about 1 order of magnitude lower than that obtained at 25 °C. This means that the lowest light intensity that can be discriminated from the background signal is about 10 times lower at 10 °C than at 25 °C. No significant improvement of the performance of the CCD sensor was observed at even lower temperatures. Therefore, in all the experiments described in the following, the temperature of the CCD sensor was set to 10 °C. Spatial Resolution. In the absence of a lens-based optical system to focus the CL signals on the CCD sensor, the maximum spatial resolution in imaging measurements would be achieved by immobilizing the biospecific capture probes directly on the sensor surface.31 However, as mentioned before, the fiber optic taper was necessary due to the low temperature of the sensor. Moreover, the development of a separate reaction chip in which the biospecific probes are immobilized onto a thin transparent support was considered much more suitable for a disposable use of the device. Therefore, the spatial resolution of the lensless CL imaging measurement was evaluated and the configuration of the reaction chip (i.e., spot/microwell sizes and separations) was optimized to minimize the cross talk between the CL signals from adjacent spots/microwells. 3181
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Figure 4. Chemiluminescence calibration curves obtained for HRP in solution in mini microtiter plates employing either CCD-based lensless imaging (0) or the LB 981 luminograph (9). Data represent the mean values ( SD obtained for three replicates.
Figure 3. (A) Chemiluminescent images and light intensity profiles of HRP spots (spot diameter ∼1 mm, spot-to-spot separation 4 mm) deposited on the glass slide obtained either with the lensless CCD imaging device (top) or with the LB 981 luminograph (bottom). (B) Cross talk (mean value ( SD evaluated along horizontal, vertical, and diagonal directions in a square array configuration, as shown in the inset) measured in inkjet printed masks for different spot sizes, i.e., 1 mm (9), 2 mm (O), and 3 mm (2), and separations. (C) Cross talk between two adjacent microwells as a function of their relative CL intensity: bars show the CL emission of the two wells, and the line plot reports the cross talk (mean values ( SD obtained for three replicates).
As expected, lensless imaging yielded lower resolution than optics-based imaging, as shown in Figure 3A in which images of CL spots acquired either with the LB 981 luminograph or with the CCD contact imaging device are compared. To evaluate different spot/microwell array geometries, black masks with square arrays of circular spots of different diameters (from 1.0 to 3.0 mm) and separations (from 1.0 to 6.0 mm) were inkjet printed on 150 μm thick clear acetate foils. The masks were then imaged with the lensless CCD device under a diffuse ambient illumination to simulate CL spots with well-defined geometric properties. For each mask, the light intensity was measured in correspondence of a spot area, as well as in three similar blind areas along the horizontal, vertical, and diagonal directions in a square array configuration (see inset in Figure 3B). Cross talk values (Figure 3B), calculated as the percentage of light crossing
from a luminous spot to adjacent blind areas, were approximately 2% or less for the geometries used in this work, i.e., mini microtiter plates (well diameter ∼2 mm, well-to-well separation 4.5 mm), flow cells (cell diameter 3.0 mm, cell-to-cell separation 6.0 mm), and arrays of spots of biospecific probes (spot diameter 1 mm, spot-to-spot separation 6.0 mm). In addition, to investigate the effect of the intensity of the CL signal on the extent of cross talk, we measured two adjacent microwells, one with a constant weak CL emission and the other with increasing CL intensities. In this case, the cross talk was evaluated as the apparent increase in the CL emission of the weakest microwell due to the light crossing from the more intense one. According to Figure 3C, the cross talk increased with the intensity of the CL emission; nevertheless, even for CL signals close to the saturation of the CCD sensor, it remained below 3%. For the sake of comparison, the data shown in Figure 3B were obtained for light intensities around 30 000 RLU. Light Detectability. The analytical performance of the CCD device for lensless imaging was assessed by producing calibration curves for HRP in solution and comparing these curves to those obtained with the reference Night Owl LB 981 luminograph (Figure 4). According to the calibration curves, the limits of detection, determined as the concentration of HRP leading to a signal corresponding to that of the blank plus three standard deviations, were 2.3 pmol/L for the CCD lensless CCD device and 1.6 pmol/L for the LB 981 luminograph. The similarity of the two values demonstrates that a good performance can be obtained with imaging even with relatively low-cost CCD sensors, which is critical in the view of the development of portable analytical devices with performance adequate to perform point-of-care analyses. To assess the increase in light collection efficiency achieved by using the lensless imaging configuration, HRP calibration curves were also obtained by employing the CCD sensor coupled with a Computar objective (8 mm, 1:1.4) obtained from CBC (AMERICA) Corp. (Commack, NY). The higher limit of detection obtained in this case (5.7 pmol/L of HRP) demonstrated that lensless imaging was more sensitive than conventional lens-based imaging, as a result of the higher light collection efficiency. On the contrary, the CMOS sensor displayed a much lower detectability, with a limit of detection of 28 pmol/L of HRP in the lensless imaging configuration. CMOS image sensors have been proposed for integrated optical detection in a variety of microfluidic systems, also in contact imaging configurations, mainly employing transmission imaging or fluorescence detection.20 For those applications, CMOS sensors offer adequate sensitivity and present several advantages, 3182
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Analytical Chemistry such as low cost, possibility to enclose sensor and control electronics in the same chip, no need for cooling, and low power consumption. However, CL emissions (despite their higher signalto-noise ratio) are usually much weaker than fluorescence, and CMOS sensors might not be sensitive enough to achieve the detection limits obtained with CCD sensors.32,33 Immobilization of Biospecific Probes. For the binding of the biospecific capture probes (antibodies and nucleic acid probes), the glass surface of the reaction chip was activated with an epoxyterminated silane derivative (GOPS) to form a self-assembled monolayer. Then biospecific capture probes, spotted in correspondence to channels, were covalently immobilized through the reaction of the GOPS epoxy groups with the free amino groups of the probes. The optimal experimental parameters for epoxy coating were evaluated by preparing arrays (8 5 spots) of HRP enzyme (1 μmol/L in 0.1 M PBS buffer) employing the procedure described in the Experimental Methods section for antibodies immobilization. The amount of HRP bound in each spot was then determined by CL imaging with the LB 981 luminograph, upon addition of the HRP CL enzyme substrate. When different immobilization conditions (i.e., GOPS concentrations ranging from pure GOPS to 0.1% v/v in toluene and incubation times from 1 to 4 h) were evaluated, it was found that the highest and most reproducible light signal intensity was obtained employing 10% v/v GOPS in toluene and 2 h incubation. A satisfactory spot-to-spot reproducibility was obtained (CV = 20%), taking into account that a manual spotting procedure was employed (data not shown). Model Application 1: Measurement of Alkaline Phosphatase Activity in Solution. The serum ALP test, selected in this work as
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an example of a CL-based enzyme activity assay in solution, is a common clinical chemistry analysis, with ALP as a biomarker of liver or bone diseases (normal reference values, which are ageand gender-related, usually range between 30 and 150 IU/L). Most of the enzymes or substrates of clinical interest can be detected using CL or bioluminescence (BL) detection, either by a single enzyme reaction or by coupling two or more enzymatic reactions.3436 Therefore, the combination of CL (or BL) enzyme or substrate assays with immunoassays and gene hybridization assays in the proposed device will certainly enlarge the range of its potential applications. Alkaline phosphatase activity was measured using the flow cell configuration of the reaction chip to increase the volume of solution under measurement for a given imaged area, thus improving assay sensitivity and reproducibility. In this work, the two flow cells in the reaction chip were exploited to perform measurements in duplicate, while they might also be used to simultaneously analyze a sample and a quality control. The CL substrate used for this measure is commonly employed in bioanalytical applications for the CL detection of ALP as an enzyme marker for biospecific probes. As reported in Figure 5A, the intensity of the CL emission was directly proportional to the ALP activity, thus allowing its quantitative measurement. The CL intensity also increased with time; therefore, the performance of the assay depends on the time after injection at which the CL signal is measured. For example, the calibration curve obtained at 3 min after injection showed a dynamic range from 10 to 230 IU/L of ALP. Light measurement can be also performed at 1 min upon injection: the dynamic range of the assay is narrower (10115 IU/L) but still adequate to discriminate normal from elevated ALP serum levels.
Figure 5. Chemiluminescence calibration curves obtained for the model assays with CCD-based lensless CL imaging detection. (A) calibration curves for ALP, obtained by measuring the CL signal at different times after injection: 1 min (b), 2 min (9), 3 min (2), and 5 min (1). (B) Calibration curve for immunoassay of HRP antigen with immobilized anti-HRP antibody. (C) Calibration curve for Parvovirus B19 DNA hybridization assay. Data represent the mean value ( SD of two (ALP activity measurements) or three (immunoassay of HRP and Parvovirus B19 DNA hybridization assay) replicates. (D) CL image (in pseudocolors) of the reaction chip performing a hybrid assay, combining ALP activity measurement (top channels), B19 DNA hybridization assay (middle channel), and HRP immunoassay (bottom channel). The configuration of the reaction chip was the same as shown in Figure 1. 3183
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Analytical Chemistry Model Application 2: Immunoassay of HRP Antigen with Immobilized Anti-HRP Antibody Microarrays. As a proof-ofconcept for immunoassay application, an assay in which HRP was captured by arrayed anti-HRP antibodies was employed as a model system. This application will mimic a conventional sandwich-type immunoassay, in which the target analyte in the sample is captured by immobilized antianalyte antibodies and revealed by a second HRP-labeled antibody upon addition of a CL substrate for HRP. To assess the optimal concentration of anti-HRP immobilized antibodies, spot arrays of antibody (ranging from 10 ng/spot to 1 μg/spot) were immobilized in different channels, and an HRP solution at the highest concentration in the calibration curve (i.e., 570 pmol/L) was flowed in the channels. Then, upon addition of the HRP CL substrate, the reaction chip was imaged to determine the amount of HRP bound to the spots. It was observed that the amount of captured HRP increased with immobilized antibody up to 0.1 μg/spot (data not shown), while no further increase was detected at higher concentration. Therefore, this amount of antibody was in excess with respect to the analyte to be captured, thus fulfilling the requirements for sandwich-type immunoassays. Successively, a calibration curve for the HRP immunoassay was produced by analyzing HRP solutions at concentrations ranging from 1.5 103 to 570 pmol/L (Figure 5B). According to this calibration curve, a limit of detection of 3.5 fmol/L (corresponding to 0.70 amol of HRP) was obtained, thus confirming the possibility to achieve adequate detection limits for conventional sandwichtype CL immunoassays using the lensless CCD imaging device. Model Application 3: Determination of Parvovirus B19 DNA with Immobilized Oligonucleotide Probe Microarrays. To verify the performance of the device for detecting nucleic acid sequences, the Parvovirus B19 was selected as a model analyte. After amplification and labeling with digoxigenin of the Parvovirus B19 target sequence, the product of amplification was denatured to obtain single strands and flowed in the channels of the reaction chip functionalized with a complementary capture probe. Hybrids were then detected by HRP-labeled antidigoxigenin antibody, CL substrate, and lensless CL imaging. As for the HRP immunoassay, preliminary experiments were carried out in order to establish the optimal amount of immobilized probe. Among the different amounts of the oligonucleotide probe (from 0.04 to 1.0 pmol/spot) tested by analyzing samples with the highest concentration in the calibration curve (i.e., 4 μmol/L of target DNA), the final CL signal increased up to 0.2 pmol/spot of probe and then remained almost constant (data not shown). Therefore, this amount of immobilized probe was assumed to be able to bind almost all the target DNA in the sample and was used to obtain the calibration curve. Such a calibration curve (shown in Figure 5C) demonstrated the ability of the system to detect amplification products with adequate sensitivity for diagnostic applications. Indeed, the target DNA could be detected down to 0.05 μmol/L, corresponding to 50 fmol of DNA amplification product. Hybrid Analytical System. To demonstrate the possibility of hybrid assays, the three model analyses described above were simultaneously performed in a reaction chip configured as shown in Figure 1 (Figure 5D). In such a way, at least three different analytes, eventually of different nature, could be determined in the same sample. However, since the number of analyses that can be carried out in a single reaction chip is limited, the approach followed for calibration and quantitative analysis should be carefully defined and optimized. With dependence on the overall
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reproducibility of the analytical procedure, a set of representative calibration curves for the various assays might be obtained upon production of each lot of reaction chips or, alternatively, a preliminary calibration using standard analyte solutions or quality control samples could be performed in a separate reaction chip before analysis of unknown samples.
’ CONCLUSIONS A new portable and versatile device suitable for “personalized diagnostics” and point-of-care multiplexed CL-based analysis has been developed exploiting a lensless imaging approach. Chemiluminescence detection was employed due to its peculiar characteristics, such as high signal detectability also in small sample volumes and large dynamic range of measurements. In addition, it provided the highest flexibility in the design of the device because the only requirement for the measurement cell is to capture as many photons as possible, which was perceived by employing this device for lensless imaging. A thermoelectrically cooled CCD sensor was employed for CL signal imaging and quantification. Sensor cooling, which led to a significant improvement in light signal detectability, does not compromise the portability of the device, which is small in size and can be battery operated. The performance of the CCD sensor was evaluated in comparison with a laboratory optic-based luminograph, showing excellent light detectability. A simple reaction chip was developed and coupled with the CCD sensor via a fiber optic taper to perform lensless imaging detection, thus obtaining a compact detection device still providing adequate imaging resolution without employment of lensbased optics. The tapered configuration of the fiber optic element also allowed enlarging the useful measurement area with respect to the size of the CCD sensor, thereby permitting image areas of the reaction chip as large as 20 15 mm2 (the reaction chip can host three separate fluidic channels, each of them containing three/four spots of immobilized biospecific probes). The device has been demonstrated to be suitable for hybrid multiplexing analysis, in which ultrasensitive immuno- and genebased assays can be performed together with simple enzyme activity measurements. The flexibility in the design of the reaction chip will permit realization of any combination of these different types of assays; therefore, reaction chips for the simultaneous detection of various biomarkers, even of different nature (enzymes, enzyme substrates, antigens, antibodies, and nucleic acids), could be easily developed. Panel assays for “personalized medicine” applications, where “ad hoc” biomarkers need to be analyzed (eventually together with conventional clinical chemistry tests and/or quantification of therapeutic drug levels) for the monitoring of the health status and the drug efficacy in a specific patient responder, represent a possible appealing application of this device. With this respect, this device could be employed by a nurse, a physician or a technician directly in the doctor’s office or at the patient’s bed to rapidly provide all the clinical chemistry information necessary for accurate diagnosis and patient follow-up. ’ AUTHOR INFORMATION Corresponding Author
*Address: Analytical and Bioanalytical Chemistry Laboratory, Department of Pharmaceutical Sciences, University of Bologna, Via Belmeloro 6, 40126 Bologna, Italy. Phone and fax: þ39 051 343398. E-mail:
[email protected]. URL: http://www.anchem.unibo.it. 3184
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