LETTERS nature PUBLISHED ONLINE: 13 DECEMBER 2009 I DOI: 10. 1038/NNANO 2009.353 nanotechnology Label-free biomarker detection from whole blood Eric Stern, Aleksandar Vacic,, Nitin K Rajan, Jason M. Criscione ', Jason Park, Bojan R Ilic, David J. Mooney, Mark A. Reed25* and Tarek M. Fahmy,6* Label-free nanosensors can detect disease markers to provide To overcome these limitations we have developed a new in-line point-of-care diagnosis that is low-cost, rapid, specific and sen- microfabricated device that operates upstream of the nanosensors sitive-1. However, detecting these biomarkers in physiological to purify biomarkers of interest. This microfluidic purification chip fluid samples is difficult because of problems such as biofouling (MPC) captures cancer biomarkers from physiological solution and non-specific binding, and the resulting need to use purified and, after washing, releases the antigens"7 into a pure buffer suitable buffers greatly reduces the clinical relevance of these sensors. for sensing. The chip design increases nanosensor specificity to that within the sensor to perform purification and detection. a antibodies to bind biomarkers for a positive signal to be produtieo Here we overcome this limitation by using distinct components of conventional sandwich techniques, because it requires tw microfluidic purification chip simultaneously captures multiple Figure 1 schematically illustrates the operation of the MPC. The biomarkers from blood samples and releases them, after avidin-functionalized chip(Fig. la)is treated with antibodies to washing, into purified buffer for sensing by a silicon nanoribbon any number of specific biomarkers conjugated to biotinylated detector. This two-stage approach isolates the detector from photocleavable crosslinkers containing a specific 19-mer dNA the complex environment of whole blood, and reduces its sequence( Fig 2a)2. The MPC geometry was chosen to optimize minimum required sensitivity by effectively pre-concentrating biomarker binding( Supplementary Fig S1) and chips were fabri- the biomarkers. We show specific and quantitative detection cated from 4-inch silicon wafers in a one-step photolithographic of two model cancer antigens from a 10 ul sample of whole process(Supplementary Fig. S2). Completed chips(Fig. 2b)were blood in less than 20 min. This study marks the first use of loaded into a custom-machined flow chamber(Fig. 2b, inset, and label-free nanosensors with physiological solutions, position- Supplementary Fig. S3), which enabled fluid handling and mair ing this technology for rapid translation to clinical settings. tained a constant volume of 5 ul in the system Biomarkers have emerged ortant diagnostic An example operation is illustrated in Fig. 1b-d. First, a blood tools for cancer and many other diseases. Continuing discoveries sample flows through the chip(Fig. 1b)and the chip-bound anti of such biomarkers and their aggregation into molecular signatures bodies bind specific soluble biomarkers, essentially purifying these suggests that multiple biomarkers will be necessary to precisely molecules from whole blood. After this capture step, wash and define disease states. Thus, parallel detection of biomarker arrays sensing buffers are perfused through the device. Flow is then is essential for translation from benchtop discovery to clinical vali- halted, and the sensing buffer-filled MPC is irradiated with ultra dation. Such a technique would enable rapid, point-of-care(POC) violet(UV) light(Fig. Ic), resulting in cleavage of the photolabile applications requiring immediate diagnosis from a physiological group20-23 and release of the bound biomarker- body-DNA sample. Critically, such a system must also be capable of detecting complexes. The UV photocleavage process was shown not to very low levels of aberrant genes and proteins, as many biomarkers affect the noactivity of the biomarkers (Supplement are present at minute concentrations during early disease phases-6. Fig. S4). The DNA component was critical for preliminary assay Given these requirements, the use of conventional diagnostic validation experiments(Fig. 2c). As shown in Fig. ld, after a says.. has been a limiting factor. An approach that is based second valve switching step transfers MPC contents to the nanosen on rapid, label-free sensing technologies would be ideally suited sor chip, the complexes bind the secondary antibodies on the nano- for clinical applications. wire surfaces. The purification/sensing operation thus requires two Since their introduction in 2001 (ref. 7), label-free nanosensors specific antibody binding events for detection, a significant improv have demonstrated great potential to serve as Poc detectors ment in selectivity over previous label-free nanosensing schemes- capable of ultrasensitive, real-time, multiplexed detection of mul- To demonstrate the effectiveness of the capture-release tiple biomolecular species68-13. Despite their appeal, electronic approach, we used a readily available fluorescently labelled nanosensors continue to be a challenge to imple because fun- antigen-antibody pair, fluorescently labelled chicken ovalbumin damental limitations render them incapable of sensing molecules in (OVA-FITC) and its antibody anti-OVA IgG. OVA-FITC w complex, physiological solutions, 8-13. Biofouling and non-specific added to heparinized murine blood and flo binding readily degrade the minute active surface areas of such OVA functionalized chip. After washo ated specific OVA-FITC devices(<0. 1 um ref. 15)and label-free sensing requires purified, buffer, fluorescence imaging demonst precisely controlled buffers to enable measurements to be per- binding to chip-bound antibodies(Fig. 2d). A control chip, formed. In the case of nanowire field-effect transistor (FET) which anti-prostate specific antigen(PSA) was bound, showed sensing, low salt (<l mM) buffers are required to prevent negligible fluorescent signal(Fig. 2d, inset). After UV irradiation reening of the charge-based electronic signal 2I and subsequent flushing of the sensing reservoir with fresh buffer, Department of Biomedical Engineering, School of Engineering and Applied Science, Yale University, New Haven, Connecticut 06511, USA, Department of Electrical Engineering, School of Engineering and Applied Science, Yale University, New Haven, Connecticut 06511, USA, Cornell Nanofabrication Facility arnell University, Ithaca, New York 14853, USA " Department of Bioengineering, School of Engineering and Applied Science, Harvard University, bridge, Massachusetts 02138, USA, Department of Applied Physics, School of Engineering and Applied Science, Yale University, New Haven Connecticut 06511, USA, Department of Chemical Engineering, School of Engineering and Applied Science, Yale University, New Haven, Connecticut 06511, USA.e-mail: mark. reed @yale. edu; tarek fahmy @yale. edu 138 NaturENanoteChnOlogyIVol5iFebRuaRy2010Iwww.nature.com/naturenanotechnology 2010 Macmillan Publishers Limited. All rights reserved
Label-free biomarker detection from whole blood Eric Stern1 , Aleksandar Vacic2, Nitin K. Rajan2, Jason M. Criscione1 , Jason Park1 , Bojan R. Ilic3, David J. Mooney4, Mark A. Reed2,5* and Tarek M. Fahmy1,6* Label-free nanosensors can detect disease markers to provide point-of-care diagnosis that is low-cost, rapid, specific and sensitive1–13. However, detecting these biomarkers in physiological fluid samples is difficult because of problems such as biofouling and non-specific binding, and the resulting need to use purified buffers greatly reduces the clinical relevance of these sensors. Here, we overcome this limitation by using distinct components within the sensor to perform purification and detection. A microfluidic purification chip simultaneously captures multiple biomarkers from blood samples and releases them, after washing, into purified buffer for sensing by a silicon nanoribbon detector. This two-stage approach isolates the detector from the complex environment of whole blood, and reduces its minimum required sensitivity by effectively pre-concentrating the biomarkers. We show specific and quantitative detection of two model cancer antigens from a 10 ml sample of whole blood in less than 20 min. This study marks the first use of label-free nanosensors with physiological solutions, positioning this technology for rapid translation to clinical settings. Biomarkers have emerged as potentially important diagnostic tools for cancer and many other diseases. Continuing discoveries of such biomarkers and their aggregation into molecular signatures suggests that multiple biomarkers will be necessary to precisely define disease states. Thus, parallel detection of biomarker arrays is essential for translation from benchtop discovery to clinical validation. Such a technique would enable rapid, point-of-care (POC) applications requiring immediate diagnosis from a physiological sample. Critically, such a system must also be capable of detecting very low levels of aberrant genes and proteins, as many biomarkers are present at minute concentrations during early disease phases3–6. Given these requirements, the use of conventional diagnostic assays5,6,14 has been a limiting factor. An approach that is based on rapid, label-free sensing technologies would be ideally suited for clinical applications6–13. Since their introduction in 2001 (ref. 7), label-free nanosensors have demonstrated great potential to serve as POC detectors capable of ultrasensitive, real-time, multiplexed detection of multiple biomolecular species6,8–13. Despite their appeal, electronic nanosensors continue to be a challenge to implement, because fundamental limitations render them incapable of sensing molecules in complex, physiological solutions6,8–13. Biofouling and non-specific binding readily degrade the minute active surface areas of such devices (,0.1 mm2 ; ref. 15) and label-free sensing requires purified, precisely controlled buffers to enable measurements to be performed. In the case of nanowire field-effect transistor (FET) sensing, low salt (, 1 mM) buffers are required to prevent screening of the charge-based electronic signal12,16. To overcome these limitations we have developed a new in-line microfabricated device that operates upstream of the nanosensors to purify biomarkers of interest. This microfluidic purification chip (MPC) captures cancer biomarkers from physiological solutions and, after washing, releases the antigens17 into a pure buffer suitable for sensing. The chip design increases nanosensor specificity to that of conventional sandwich assay techniques, because it requires two antibodies to bind biomarkers for a positive signal to be produced18. Figure 1 schematically illustrates the operation of the MPC. The avidin-functionalized chip19 (Fig. 1a) is treated with antibodies to any number of specific biomarkers conjugated to biotinylated, photocleavable crosslinkers containing a specific 19-mer DNA sequence (Fig. 2a)20. The MPC geometry was chosen to optimize biomarker binding (Supplementary Fig. S1)14 and chips were fabricated from 4-inch silicon wafers in a one-step photolithographic process (Supplementary Fig. S2). Completed chips (Fig. 2b) were loaded into a custom-machined flow chamber (Fig. 2b, inset, and Supplementary Fig. S3), which enabled fluid handling and maintained a constant volume of 5 ml in the system. An example operation is illustrated in Fig. 1b–d. First, a blood sample flows through the chip (Fig. 1b) and the chip-bound antibodies bind specific soluble biomarkers, essentially purifying these molecules from whole blood. After this capture step, wash and sensing buffers are perfused through the device. Flow is then halted, and the sensing buffer-filled MPC is irradiated with ultraviolet (UV) light (Fig. 1c), resulting in cleavage of the photolabile group20223 and release of the bound biomarker–antibody–DNA complexes. The UV photocleavage process was shown not to affect the immunoactivity of the biomarkers (Supplementary Fig. S4). The DNA component was critical for preliminary assay validation experiments (Fig. 2c). As shown in Fig. 1d, after a second valve switching step transfers MPC contents to the nanosensor chip, the complexes bind the secondary antibodies on the nanowire surfaces. The purification/sensing operation thus requires two specific antibody binding events for detection, a significant improvement in selectivity over previous label-free nanosensing schemes6–13. To demonstrate the effectiveness of the capture–release approach, we used a readily available fluorescently labelled antigen–antibody pair, fluorescently labelled chicken ovalbumin (OVA–FITC) and its antibody anti-OVA IgG. OVA–FITC was added to heparinized murine blood and flowed through an antiOVA functionalized chip. After washing and flushing with sensing buffer, fluorescence imaging demonstrated specific OVA–FITC binding to chip-bound antibodies (Fig. 2d). A control chip, to which anti-prostate specific antigen (PSA) was bound, showed a negligible fluorescent signal (Fig. 2d, inset). After UV irradiation and subsequent flushing of the sensing reservoir with fresh buffer, 1 Department of Biomedical Engineering, School of Engineering and Applied Science, Yale University, New Haven, Connecticut 06511, USA, 2 Department of Electrical Engineering, School of Engineering and Applied Science, Yale University, New Haven, Connecticut 06511, USA, 3 Cornell Nanofabrication Facility, Cornell University, Ithaca, New York 14853, USA, 4 Department of Bioengineering, School of Engineering and Applied Science, Harvard University, Cambridge, Massachusetts 02138, USA, 5 Department of Applied Physics, School of Engineering and Applied Science, Yale University, New Haven, Connecticut 06511, USA, 6 Department of Chemical Engineering, School of Engineering and Applied Science, Yale University, New Haven, Connecticut 06511, USA. *e-mail: mark.reed@yale.edu; tarek.fahmy@yale.edu LETTERS PUBLISHED ONLINE: 13 DECEMBER 2009 | DOI: 10.1038/NNANO.2009.353 138 NATURE NANOTECHNOLOGY | VOL 5 | FEBRUARY 2010 | www.nature.com/naturenanotechnology © 2010 Macmillan Publishers Limited. All rights reserved.
NATURE NANOTECHNOLOGY DOl: 10.1038/NNANO 2009.353 LETTERS □ Nanosensors Photocleavable □ 1 anti-PSA □2 anti-CA153 cA15.3 口 Spiked bl photocleavable crosslinker to the MPC. The chip is placed in a plastic housing and a valve(pink) directs fluid flow exiting the chip to either a waste th a Figure 1 I Schematic of MPC operation. a, Primary antibodies to multiple biomarkers, here PSa and carbohydrate antigen 15.3(CA15 3), are bound wi eptade or the nanosensor chip. b, Whole blood is injected into the ith the valve set to the waste compartment (black arrow shows the direction of luid flow) and, if present in the sample, biomarkers bind their cognate antibodies. c, Washing steps follow blood flow, and the chip volume(5 ul) is filled with sensing buffer before UV irradiation (orange arrows). During UV exposure the photolabile crosslinker cleaves, releasing the antibody-antigen complexes into solution. d, The valve is set to the nanosensor reservoir(black arrow shows the direction of fluid flow) and the s ul volume is transferred, enabling label free sensing to be performed to determine the presence of specific biomarkers the fluorescence signal from the anti-OVA chip was greatly dimin- and microscale lateral dimensions2, which are less sensitive but ished(Fig. 2d, inset) have significant fabrication and cost advantages. These devices, fab To demonstrate the generality of the MPC technique, we used ricated using conventional lithographic techniques, have b two model cancer antigens, PSA and carbohydrate antigen 15.3 demonstrated to detect streptavidin in the 0.0318-53 ng m ( CA153), standard clinical markers for prostate2425 and breast range2, a sensitivity range ideally suited for MPC-purified cancer ancer26.27, respectively. Successful capture and release of PSA and antigen detection. We fabricated 25-nm-thick devices according to CA153 was verified with a modified enzyme-linked immunoassay a similar process( Supplementary Fig. S5)28 but incorporated (ELISA)technique(Fig. 2c)8, in which the first detection step con- ohmic contacts to the devices. Device images are provided in sisted of the hybridization of a complementary, biotinylated 19-mer Fig. 3a. Electrical characterization verified that this approach pro- to the crosslinker DNA sequence. Six increasing concentrations of duced high-quality devices, with on/off ratios of >106(Fig. 3b) PSA and CA153 were added to heparinized rat blood and and small hystereses between forward and reverse IDs(VG) sweeps panned clinically relevant ranges24-77 The data in Fig. 2e, f demon- electrical characteristics(Supplementary Fig. S7), and solution trate a monotonic relationship between the concentration of bio- gating(VG soln) demonstrated that VG=-5V was an optim marker introduced in whole blood and that released into pure operating point for sensing studies(Fig. 3d). sensing buffer. The absolute yields of these experiments are in agree- As shown in Fig. 2 and detailed in the Supplementary ment with modelling studies(Supplementary Fig. SIc). Biomarker Information, devices were functionalized either with anti-PSA or capture by MPCs can be significantly increased by adjusting anti-CA15.3. Antibodies were immobilized to the sensor using either the operation conditions, such as the flow rate into the N-hydroxysuccinimide(NHS)/1-ethyl-3-(3-dimethylaminopropyl) device (modelled in Supplementary Fig. Sld), or the carbodiimide(EDC)chemistry. To verify that the signal from device dimensions A critical feature of this integrated approach is that the MPC- direct measurements of the amount of the signal that would be purified biomarker complex concentrations are well above those unscreened were carried out by varying buffer salt concentration required for label-free, electronic detection. Although previous This study indicated that 50% of the signal was not screened by studies using nanowire sensors have demonstrated PSA detection the buffer solution(see Supplementary Information) levels as low as 0.9 pg ml(refs 6, 10), this exquisite sensitivity is We applied these devices to sensing the biomarkers from the a critical factor for MPC-nanosensor operation. We thus MPC-purified whole blood samples. The normalized res ose to use ' nanoribbons, devices with nanoscale thicknesses these same devices to MPC-purified, antigen-spiked blood NaturENanoteChnOlogyIVol5iFebRuaRy2010Iwww.nature.com/naturenanotechnology 139 @2010 Macmillan Publishers Limited. All rights reserved
the fluorescence signal from the anti-OVA chip was greatly diminished (Fig. 2d, inset). To demonstrate the generality of the MPC technique, we used two model cancer antigens, PSA and carbohydrate antigen 15.3 (CA15.3), standard clinical markers for prostate24,25 and breast cancer26,27, respectively. Successful capture and release of PSA and CA15.3 was verified with a modified enzyme-linked immunoassay (ELISA) technique (Fig. 2c)18, in which the first detection step consisted of the hybridization of a complementary, biotinylated 19-mer to the crosslinker DNA sequence. Six increasing concentrations of PSA and CA15.3 were added to heparinized rat blood and samples were flowed through MPCs functionalized with both anti-PSA and anti-CA15.3. The introduced concentrations spanned clinically relevant ranges24–27. The data in Fig. 2e,f demonstrate a monotonic relationship between the concentration of biomarker introduced in whole blood and that released into pure sensing buffer. The absolute yields of these experiments are in agreement with modelling studies (Supplementary Fig. S1c). Biomarker capture by MPCs can be significantly increased by adjusting either the operation conditions, such as the flow rate into the device (modelled in Supplementary Fig. S1d), or the device dimensions. A critical feature of this integrated approach is that the MPCpurified biomarker complex concentrations are well above those required for label-free, electronic detection. Although previous studies using nanowire sensors have demonstrated PSA detection levels as low as 0.9 pg ml21 (refs 6,10), this exquisite sensitivity is not a critical factor for MPC-nanosensor operation. We thus chose to use ‘nanoribbons’, devices with nanoscale thicknesses and microscale lateral dimensions28, which are less sensitive but have significant fabrication and cost advantages. These devices, fabricated using conventional lithographic techniques, have been demonstrated to detect streptavidin in the 0.0318–53 ng ml21 range28, a sensitivity range ideally suited for MPC-purified cancer antigen detection. We fabricated 25-nm-thick devices according to a similar process (Supplementary Fig. S5)28, but incorporated ohmic contacts to the devices. Device images are provided in Fig. 3a. Electrical characterization verified that this approach produced high-quality devices, with on/off ratios of .106 (Fig. 3b) and small hystereses between forward and reverse IDS(VG) sweeps (Fig. 3c), where IDS is the drain–source current and VG is the gate voltage. Surface functionalization did not compromise the device electrical characteristics (Supplementary Fig. S7), and solution gating (VG,SOLN) demonstrated that VG ¼ 2 5 V was an optimal operating point for sensing studies (Fig. 3d). As shown in Fig. 2 and detailed in the Supplementary Information, devices were functionalized either with anti-PSA or anti-CA15.3. Antibodies were immobilized to the sensor using N-hydroxysuccinimide (NHS)/1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) chemistry. To verify that the signal from binding proteins would not be screened by the buffer solution, direct measurements of the amount of the signal that would be unscreened were carried out by varying buffer salt concentration16. This study indicated that 50% of the signal was not screened by the buffer solution (see Supplementary Information). We applied these devices to sensing the biomarkers from the MPC-purified whole blood samples. The normalized responses of these same devices to MPC-purified, antigen-spiked blood Cap-rel chip a b c d Housing/tubing Nanosensors Valve Photocleavable 1’ anti-PSA 1’ anti-CA15.3 2’ anti-PSA 2’ anti-CA15.3 PSA CA15.3 Spiked blood Buffer Figure 1 | Schematic of MPC operation. a, Primary antibodies to multiple biomarkers, here PSA and carbohydrate antigen 15.3 (CA15.3), are bound with a photocleavable crosslinker to the MPC. The chip is placed in a plastic housing and a valve (pink) directs fluid flow exiting the chip to either a waste receptacle or the nanosensor chip. b, Whole blood is injected into the chip with the valve set to the waste compartment (black arrow shows the direction of fluid flow) and, if present in the sample, biomarkers bind their cognate antibodies. c, Washing steps follow blood flow, and the chip volume (5 ml) is filled with sensing buffer before UV irradiation (orange arrows). During UV exposure, the photolabile crosslinker cleaves, releasing the antibody–antigen complexes into solution. d, The valve is set to the nanosensor reservoir (black arrow shows the direction of fluid flow) and the 5 ml volume is transferred, enabling labelfree sensing to be performed to determine the presence of specific biomarkers. NATURE NANOTECHNOLOGY DOI: 10.1038/NNANO.2009.353 LETTERS NATURE NANOTECHNOLOGY | VOL 5 | FEBRUARY 2010 | www.nature.com/naturenanotechnology 139 © 2010 Macmillan Publishers Limited. All rights reserved.
LETTERS NATURE NANOTECHNOLOGY DOL: 10.1038/NNANO. 2009.353 PSA detectio DNA bound to 1 antibody 种A oncentration PSA introduced (ng ml-) Concentration CA153 introduced (units ml-y) Figure 2 I MPC operation. a, Molecular structure of the photocleavable crosslinker. Primary antibody conjugation was performed with the amino group(right) and binding to chip -bound avidin occurred through the biotin group (left). b, Scanning electron micrograph of a representative(w=4 mm)x (=7 mm)x(h=100 um) MPC capture-release chip. The inset is an optical image of MPC operation during washing. c, Schematic representation of PSA and CA153 detection using a modified ELISA technique. d, Fluorescence optical micrograph of an anti-oVA functionalized MPC following OVA-FITC-spiked whole blood flow and washing. The inset plots the pixel intensity (grey value, determined by Image ) versus position for the red cut line (green data plot and similar cut lines from images of post-UV irradiation and transfer(blue) and of an anti-PSA functionalized MPC following OVA-FITc-spiked blood flow and washing. The same exposure times were used for all images. e, f Scatter plots showing the concentration of PSa(e) and CA153 (f)released from the MPC versus the concentration of PSA and CA153 introduced in whole blood, respectively. Each data point represents the average of three separate MPC runs, and error bars represent one standard deviation. samples containing both 2.5 ng ml-I PSA and 30 U ml-I CA153 interactions. Thus, we focus on the initial kinetic reaction rates (as well as negative controls) are shown in Fig. 4a, b, respectively. instead of endpoint detection. After the injection transient noise subsided, device current levels Using these rates, a quantification of analyte concentrations were increased by antigen binding due to the negative charge con-(against a known) can be made, as shown in Fig. 4c, d. whole ferred to the antigens by the basic sensing buffer. Similar signals blood samples spiked with 2 ng ml PSA and 15 U ml CA153 were obtained with a PSA/CA153 spiked sensing buffer positive were MPC purified and sensed with anti-PSA and anti-CA153 control, and no device response was observed with an unspiked, functionalized devices. Using the slope of the normalized device MPC-purified blood negative control. To reduce potential transient temporal response, we find that the slope ratios of both the PSA electrical signals upon injection, buffer salt concentrations of the and CA153 responses agree quite well with the initial spiked functionalized devices and the MPC-purified samples were kept whole biomarker concentrations. For PSA, the slope ratio is 1.38 approximately the same. The positive signal was observed to compared with a concentration ratio of 1.25; for CA15 3, the slope ncrease linearly with time, following well-known ligand-receptor ratio is 1.94, compared with a concentration ratio of 2. 0. It should kinetics, in which initial rates at low relative analyte concentrations be noted that this quantification occurs in the presence of another are directly proportional to species concentration. In fact, the species, therefore also demonstrating selectivity(see Supplementary asymptotic saturation value of the device response is weakly Information for further repeatability dat dependent on concentration for reversible reactions with a low dis- The integration of a microfluidic purification step with label-free sociation constant2, which is the case for the antigen-antibody nanosensor detection represents a paradigm shift in label-free 140 NaturENanoteChnOlogyIVol5iFebRuaRy2010Iwww.nature.com/naturenanotechnology 2010 Macmillan Publishers Limited. All rights reserved
samples containing both 2.5 ng ml21 PSA and 30 U ml21 CA15.3 (as well as negative controls) are shown in Fig. 4a,b, respectively. After the injection transient noise subsided11, device current levels were increased by antigen binding due to the negative charge conferred to the antigens by the basic sensing buffer. Similar signals were obtained with a PSA/CA15.3 spiked sensing buffer positive control, and no device response was observed with an unspiked, MPC-purified blood negative control. To reduce potential transient electrical signals upon injection, buffer salt concentrations of the functionalized devices and the MPC-purified samples were kept approximately the same. The positive signal was observed to increase linearly with time, following well-known ligand–receptor kinetics29, in which initial rates at low relative analyte concentrations are directly proportional to species concentration30. In fact, the asymptotic saturation value of the device response is weakly dependent on concentration for reversible reactions with a low dissociation constant29, which is the case for the antigen–antibody interactions. Thus, we focus on the initial kinetic reaction rates instead of endpoint detection30. Using these rates, a quantification of analyte concentrations (against a known) can be made, as shown in Fig. 4c,d. Whole blood samples spiked with 2 ng ml21 PSA and 15 U ml21 CA15.3 were MPC purified and sensed with anti-PSA and anti-CA15.3 functionalized devices. Using the slope of the normalized device temporal response, we find that the slope ratios of both the PSA and CA15.3 responses agree quite well with the initial spiked whole biomarker concentrations. For PSA, the slope ratio is 1.38, compared with a concentration ratio of 1.25; for CA15.3, the slope ratio is 1.94, compared with a concentration ratio of 2.0. It should be noted that this quantification occurs in the presence of another species, therefore also demonstrating selectivity (see Supplementary Information for further repeatability data). The integration of a microfluidic purification step with label-free nanosensor detection represents a paradigm shift in label-free S H N O O N H NO2 DNA 19-mer HN NH O NH2 1.8 PSA detection a c d e f b CA15.3 detection Streptavidin-HRP complex DNA bound to 1’ antibody Complementary DNA bound to biotin 1.6 1.4 1.2 1.0 0.8 0.6 0.4 0.2 0.0 0.0 0.5 1.0 1.5 2.0 2.5 3.0 3.5 4.0 4.5 1 2 3 4 5 6 7 8 9 10 10 20 30 40 50 Concentration PSA introduced (ng ml−1) Concentration CA15.3 introduced (units ml−1) Concentration PSA captured and released (ng ml−1) Concentration CA15.3 captured and released (units ml−1) 11 10 20 30 40 50 Intensity (a.u.) Position along red line Figure 2 | MPC operation. a, Molecular structure of the photocleavable crosslinker. Primary antibody conjugation was performed with the amino group (right) and binding to chip-bound avidin occurred through the biotin group (left). b, Scanning electron micrograph of a representative (w ¼ 4 mm) (l ¼ 7 mm) (h ¼ 100 mm) MPC capture–release chip. The inset is an optical image of MPC operation during washing. c, Schematic representation of PSA and CA15.3 detection using a modified ELISA technique. d, Fluorescence optical micrograph of an anti-OVA functionalized MPC following OVA–FITC-spiked whole blood flow and washing. The inset plots the pixel intensity (grey value, determined by ImageJ) versus position for the red cut line (green data plot) and similar cut lines from images of post-UV irradiation and transfer (blue) and of an anti-PSA functionalized MPC following OVA–FITC-spiked blood flow and washing. The same exposure times were used for all images. e,f, Scatter plots showing the concentration of PSA (e) and CA15.3 (f) released from the MPC versus the concentration of PSA and CA15.3 introduced in whole blood, respectively. Each data point represents the average of three separate MPC runs, and error bars represent one standard deviation. LETTERS NATURE NANOTECHNOLOGY DOI: 10.1038/NNANO.2009.353 140 NATURE NANOTECHNOLOGY | VOL 5 | FEBRUARY 2010 | www.nature.com/naturenanotechnology © 2010 Macmillan Publishers Limited. All rights reserved.
NATURE NANOTECHNOLOGY DOl: 10.1038/NNANO 2009.353 LETTERS b1x10-5 e99999 Q99ee223 1×10-80 e99 Forward Reverse 1x10-8 1×10 1×10-9 1×10-10 x101 Figure 3 Nanosensor electrical characteristics. a, Optical image of devices fitted with sensing reservoirs. The inset shows an optical micrograph of a omplete device. Only the central region of the device(black arrow) is exposed to the solution. Metal leads contact the device source and drain and fan ou to larger contacts (not shown). The 25-nm-thick silicon device appears yellow b, Ips(Vps) plot with VG varied from0 to-20 V(black arrow shows the direction of increasing negative Vo) for a representative device illustrating p-type accumulation mode behaviour. Vps drain-source voltage. c Ios(VG) plot (VDs=1 V) for the device used in b. The inset highlights Ips (nA) around the operating point(VG =-5V) d, Plot demonstrating the effect of a varying solution gate voltage(VGSoLN) on device current (ps black solid line)and device-to-solution leakage current (lEAk blue dashed line) for Vps=1 electronic sensing of biomolecules. The technique described here 3. Etzioni, R et al. The case for early detection. Nature Rev. Cancer 3, enables biomarker detection from whole blood or any other physio- 243-252(2003) logical fluid without the challenges associated with tailoring sensor 4. Liang, S.& Chan, D w. Enzymes and related proteins as cancer biomarkers. operation for the medium of interest or engineering nanosensors a proteomic approach. Clin. Chim. Acta 381, 93-97(2007). that can withstand complex fluid media. Furthermore, the need 5. Fan, R et al. Integrated barcode chips for rapid, multiplexed analysis of proteins in microliter quantities of blood. Nature BiotechnoL. 26, for ultrasensitivity in electronic detection may not be essential with such an integrated platform because of its ability to pre-con- 6. Zheng, G, Patolsky, E, Cui, Y.Wang, W U& Lieber, C.M. Multiplexed centrate molecules of choice before sensing. The attractiveness of electrical detection of cancer markers with nanowire sensor arrays. Nature the method lies in its simplicity, speed and ability to simultaneously Biotechnol23,1294-1301(2005) capture multiple biomarkers, enabling multiplexed, highly sensitive 7. Cui, Y, Wei, Q, Park, H. Lieber, C. M. Nanowire nanosensors for highly downstream detection with label-free sensors. This proof-of-prin sensitive and selective detection of biological and chemical species. Science 293, 1289-1292(2001) ciple demonstration of the non-integrated individual components 8. Jain, K. K. Nanotechnology in clinical laboratory diagnostics. Clin. Chim.Acta should be easily integratable into a compact, self-contained 358,37-54(2005) tem. Furthermore, the low cost of MPC purification rend 9. Burg. The m capable of stand-alone use or use in tandem with nanoparticles in fluid Nature 446, 1066-1069(2007) more expensive sensing methodologies, such as rare circulating 10. Kim, A et al. Ultrasensitive, label-free and real-time immunodetection using tumour cell detectors, for more complex diagnoses. The portabil- 11. Stern, E et a. abel-free immunodetection with CMoS-compatible ep semiconducting nanowires. Nature 445, 519-522(2007). label-free sensors and should position these and similar nascent 12. Stern, E, Vacic, A& Reed, M A Semiconducting nanowire sensing technologies for rapid molecular signature determinations. field-effect transistor biomolecular sensors. IEEE Trans. Electron. Dev 3119-3130(2008) Received 11 May 2009; accepted 15 October 2009, 13. Bunimovich, Y L et al. Quantitative real-time measurements of DNA published online 13 December 2009 th alkylated nonoxidized silicon nanowires in electrolyt solution. J. Am. Chem. Soc. 128, 16323-16331(2006) References 1. Sander. C. Genomic medicine and the future of health care. Science 287 microchip technology. Nature 450, 1235-1239(2007) 1977-1978(2000) 15. Gupta, A. K ef al. Anomalous resonance in a nanomechanical biosensor 2. Jemal, A. et al. Cancer statistics 2008. CA Cancer J. Clin. 58, 71-96(2008) Proc. Natl Acad. Sci. USA 103, 13362-13367(2006) NaturENanoteChnOlogyIVol5iFebRuaRy2010Iwww.nature.com/naturenanotechnology @2010 Macmillan Publishers Limited. All rights reserved
electronic sensing of biomolecules. The technique described here enables biomarker detection from whole blood or any other physiological fluid without the challenges associated with tailoring sensor operation for the medium of interest or engineering nanosensors that can withstand complex fluid media. Furthermore, the need for ultrasensitivity in electronic detection may not be essential with such an integrated platform because of its ability to pre-concentrate molecules of choice before sensing. The attractiveness of the method lies in its simplicity, speed and ability to simultaneously capture multiple biomarkers, enabling multiplexed, highly sensitive downstream detection with label-free sensors. This proof-of-principle demonstration of the non-integrated individual components should be easily integratable into a compact, self-contained system. Furthermore, the low cost of MPC purification renders this system capable of stand-alone use or use in tandem with more expensive sensing methodologies, such as rare circulating tumour cell detectors14, for more complex diagnoses. The portability and versatility of this method represents the crucial next step for label-free sensors and should position these and similar nascent sensing technologies for rapid molecular signature determinations. Received 11 May 2009; accepted 15 October 2009; published online 13 December 2009 References 1. Sander, C. Genomic medicine and the future of health care. Science 287, 1977–1978 (2000). 2. Jemal, A. et al. Cancer statistics 2008. CA Cancer J. Clin. 58, 71–96 (2008). 3. Etzioni, R. et al. The case for early detection. Nature Rev. Cancer 3, 243–252 (2003). 4. Liang, S. & Chan, D. W. Enzymes and related proteins as cancer biomarkers: a proteomic approach. Clin. Chim. Acta 381, 93–97 (2007). 5. Fan, R. et al. Integrated barcode chips for rapid, multiplexed analysis of proteins in microliter quantities of blood. Nature Biotechnol. 26, 1373–1378 (2008). 6. Zheng, G., Patolsky, F., Cui, Y., Wang, W. U. & Lieber, C. M. Multiplexed electrical detection of cancer markers with nanowire sensor arrays. Nature Biotechnol. 23, 1294–1301 (2005). 7. Cui, Y., Wei, Q., Park, H. & Lieber, C. M. Nanowire nanosensors for highly sensitive and selective detection of biological and chemical species. Science 293, 1289–1292 (2001). 8. Jain, K. K. Nanotechnology in clinical laboratory diagnostics. Clin. Chim. Acta 358, 37–54 (2005). 9. Burg, Thomas T. P. et al. Weighing of biomolecules, single cells and single nanoparticles in fluid. Nature 446, 1066–1069 (2007). 10. Kim, A. et al. Ultrasensitive, label-free and real-time immunodetection using silicon field-effect transistors. Appl. Phys. Lett. 91, 103901 (2007). 11. Stern, E. et al. Label-free immunodetection with CMOS-compatible semiconducting nanowires. Nature 445, 519–522 (2007). 12. Stern, E., Vacic, A. & Reed, M. A. Semiconducting nanowire field-effect transistor biomolecular sensors. IEEE Trans. Electron. Dev. 55, 3119–3130 (2008). 13. Bunimovich, Y. L. et al. Quantitative real-time measurements of DNA hybridization with alkylated nonoxidized silicon nanowires in electrolyte solution. J. Am. Chem. Soc. 128, 16323–16331 (2006). 14. Nagrath, S. et al. Isolation of rare circulating tumor cells in cancer patients by microchip technology. Nature 450, 1235–1239 (2007). 15. Gupta, A. K. et al. Anomalous resonance in a nanomechanical biosensor. Proc. Natl Acad. Sci. USA 103, 13362–13367 (2006). a b 1×10−5 c d 1×10−6 1×10−7 1×10−8 1×10−9 1×10−10 1×10−11 1×10−12 1×10−13 1×10−5 1×10−6 1×10−7 1×10−8 1×10−9 1×10−10 1×10−5 1×10−6 1×10−7 1×10−8 1×10−9 1×10−10 1×10−11 1×10−12 1×10−13 1×10−14 0.0 0 −5 0 −5 −10 −15 −20 −4.8 −5.0 −5.2 80 100 120 Forward Reverse −10 −15 −20 0.5 1.0 VDS (V) IDS (A) IDS ILEAK IDS (A) I(A) VG (V) VG,SOLN (V) 1.5 2.0 VG Figure 3 | Nanosensor electrical characteristics. a, Optical image of devices fitted with sensing reservoirs. The inset shows an optical micrograph of a complete device. Only the central region of the device (black arrow) is exposed to the solution. Metal leads contact the device source and drain and fan out to larger contacts (not shown). The 25-nm-thick silicon device appears yellow. b, IDS(VDS) plot with VG varied from 0 to –20 V (black arrow shows the direction of increasing negative VG) for a representative device illustrating p-type accumulation mode behaviour. VDS, drain–source voltage. c, IDS(VG) plot (VDS ¼ 1 V) for the device used in b. The inset highlights IDS (nA) around the operating point (VG ¼ –5 V). d, Plot demonstrating the effect of a varying solution gate voltage (VG,SOLN) on device current (IDS, black solid line) and device-to-solution leakage current (ILEAK; blue dashed line) for VDS ¼ 1 V. NATURE NANOTECHNOLOGY DOI: 10.1038/NNANO.2009.353 LETTERS NATURE NANOTECHNOLOGY | VOL 5 | FEBRUARY 2010 | www.nature.com/naturenanotechnology 141 © 2010 Macmillan Publishers Limited. All rights reserved
LETTERS NATURE NANOTECHNOLOGY DOL: 10.1038/NNANO. 2009.353 a1020 b105 PSA/CA153 spiked blood o PSA/CA153 spiked blood 1.015 CA15.3 -25 me(s) Time(s) +2.0 ng ml-I 15U ml-1C 1.000 Slope ratio=1.38 sAratiA-1 -250 Time(s) Figure 4 I Label-free sensing. All sensing measurements were performed at Vps=1V and VG=-5V and all sample introductions occurred at t=o Normalizations were performed by dividing device currents by the pre-addition(t <O)current level average. Vos drain-source voltage. a, Response of an anti-PSA functionalized sensor to a MPC purified blood sample initially containing 2.5 ng ml-I PSA(and also 30 U ml-l CA153), or a control sample containing neither. b, Response of an anti-CA15 functionalized sensor to a MPC-purified blood sample initially containing 30 U ml-CA153(and also 2.5 ng ml- PSA), or a control sample containing neither. c, d, Normalized response of two anti-PSA(c)and two anti-CA153(d) functionalized devices to MPC-purified blood containing both PSA and CA153, with concentrations labelled. A least-squares fit is represented by a solid black line over the selected endpoints). The ratio of the normalized slopes calibrates the ratio of concentrations. ansistor sensors Nano af the Debye screening length on nanowire field effect 29. Cantor, C.R. Schimmel, P. R. Biophysical Chemistry: Part l: The Behavior of t.7,3405-3409(2007) Biological Macromolecules( Freeman, 1980). 17. Zhou, H Ranish, J. A, Watts, J. D. Aebersold, R. Quantitative proteome 30. Homola. ]. Present and future of surface plasmon resonance biosensors. Anal Bioanal. Chen. 377, 528-539(2003). Biotechnol20,512-515(2002) 中 Comb. Chem. High Thc knows wogeieto thank L Strict for many helpful discussions.m. look 19. Hermanson, G. T. Bioconjugate Techniques(Elsevier, 1996). J. Bertram for blood samples, M. Power for device processing assistance, M. Salt: for departmental support, and D. Stern and K. Milnamow for critical reading of the 20. Bai, X, Kim, S, Li, Z Turro, N. J.& Ju, J. Design and synthesis of a photocleavable biotinylated nucleotide for DNA analysis by mass spectrometry. manuscript. The work was supported in part by the National Institute of Health(NIH) Nucleic Acids Res 32, 535-541(2004) rough grant no. ROlEBO08260(M.A R and T.M.F. )Can nstitute for Advanced Research(CIfAR), and Army Research Office(ARO)(W91INF-08-1-0365). This work was 21. Handwerger, R G. Diamond, S L Biotinylated photocleavable lyethylenimine: capture and triggered release of nucleic acids from solid ormed in part at the Cornell Nanoscale Science and Technology Facility, a member c supports. Bioconjug. Chem. 18, 717-723(2007) the National Nanotechnology Infrastructure Network that is supported by the National 23. Olejnik, I et al. Photocleavable biotin derivatives-a versatile approach for the Author contributions ation of biomolecules. Proc. Natl Acad. Sci. USA 92, 7590-7594(1995) E.S. designed the MPC and performed all MPC experiments. E.S. and B RL designed the 24. Vickers, A J, Savage, C, OBrien, M. F.& Lilja, H. Systematic review of MPC fabrication and performed MPC velocity and doubling time as predictors nanosensor fabrication process and E.S.A. V. and BRI performed nanose processing D.J.M. assisted with MPC and experimental design, and data 25. Shariat,S.F,Scardino, P.T.& Lilja, H Screening for prostate cancer. an update. analysis E.S., A V. NKR and J.M.C. performed the sensing measurements E.S. Can.. UroL15,4363-4374(2008) L.M. C and J.P. prepared and analysed the protein samples. E.S., M.A.R. and. ME. wrote 6. Rubach, M. Szymendera, J. J, Kaminska, J. Kowalska, M. Serum the manuscript and edited it, with contributions from all authors. CA 15.3, CEA and ESR patterns in breast cancer. Int. J. Biol. Markers 12, 68-173(1997) Additional information 27. Uehara, M. et aL. Long-term prog study of carcinoembryonic The authors deare competing financial interests: details accompany the paper at ww (CEA)and carbohydrate antigen 15-3(CA 15-3)in breast cancer. Inf. J Clin. Onco.13,47-451(2008) rints and permission infor Ifstrom,N.Karlstrom,A.e.&Linnros,J.Siliconnanoribbonsforelectricalonlineathttp://npgnaturecom/reprintsandpermissions/.Correspondenceandrequestsfor detection of biomolecules. Nano Letf. 8, 945-949(2008) materials should be addressed to M.A.R. and T M.F NaturENanoteChnOlogyIVol5iFebRuaRy2010Iwww.nature.com/naturenanotechnology 2010 Macmillan Publishers Limited. All rights reserved
16. Stern, E. et al. Importance of the Debye screening length on nanowire field effect transistor sensors. Nano Lett. 7, 3405–3409 (2007). 17. Zhou, H., Ranish, J. A., Watts, J. D. & Aebersold, R. Quantitative proteome analysis by solid-phase isotope tagging and mass spectrometry. Nature Biotechnol. 20, 512–515 (2002). 18. Templin, M. F., Stoll, D., Bachmann, J. & Joos, T. O. Protein microarrays and multiplexed sandwich immunoassays: what beats the beads? Comb. Chem. High Through. Screen 7, 223–229 (2004). 19. Hermanson, G. T. Bioconjugate Techniques (Elsevier, 1996). 20. Bai, X., Kim, S., Li, Z., Turro, N. J. & Ju, J. Design and synthesis of a photocleavable biotinylated nucleotide for DNA analysis by mass spectrometry. Nucleic Acids Res. 32, 535–541 (2004). 21. Handwerger, R. G. & Diamond, S. L. Biotinylated photocleavable polyethylenimine: capture and triggered release of nucleic acids from solid supports. Bioconjug. Chem. 18, 717–723 (2007). 22. Senter, P. D. et al. Novel photocleavable protein crosslinking reagents and their use in the preparation of antibody–toxin conjugates. Photochem. Photobiol. 42, 231–237 (1985). 23. Olejnik, J. et al. Photocleavable biotin derivatives—a versatile approach for the isolation of biomolecules. Proc. Natl Acad. Sci. USA 92, 7590–7594 (1995). 24. Vickers, A. J., Savage, C., O’Brien, M. F. & Lilja, H. Systematic review of pretreatment prostate-specific antigen velocity and doubling time as predictors for prostate cancer. J. Clin. Oncol. 27, 398–403 (2009). 25. Shariat, S. F., Scardino, P. T. & Lilja, H. Screening for prostate cancer: an update. Can. J. Urol. 15, 4363–4374 (2008). 26. Rubach, M., Szymendera, J. J., Kaminska, J. & Kowalska, M. Serum CA 15.3, CEA and ESR patterns in breast cancer. Int. J. Biol. Markers 12, 168–173 (1997). 27. Uehara, M. et al. Long-term prognostic study of carcinoembryonic antigen (CEA) and carbohydrate antigen 15-3 (CA 15-3) in breast cancer. Int. J. Clin. Oncol. 13, 447–451 (2008). 28. Elfstrom, N., Karlstrom, A. E. & Linnros, J. Silicon nanoribbons for electrical detection of biomolecules. Nano Lett. 8, 945–949 (2008). 29. Cantor, C. R. & Schimmel, P. R. Biophysical Chemistry: Part III: The Behavior of Biological Macromolecules (Freeman, 1980). 30. Homola, J. Present and future of surface plasmon resonance biosensors. Anal. Bioanal. Chem. 377, 528–539 (2003). Acknowledgements The authors would like to thank J. Straight for many helpful discussions, M. Look and J. Bertram for blood samples, M. Power for device processing assistance, M. Saltzman for departmental support, and D. Stern and K. Milnamow for critical reading of the manuscript. The work was supported in part by the National Institute of Health (NIH) through grant no. R01EB008260 (M.A.R. and T.M.F.), Canadian Institute for Advanced Research (CIfAR), and Army Research Office (ARO) (W911NF-08-1-0365). This work was performed in part at the Cornell Nanoscale Science and Technology Facility, a member of the National Nanotechnology Infrastructure Network that is supported by the National Science Foundation (NSF), and at the Yale Institute for Nanoscience and Quantum Engineering. This paper is dedicated to the memory of Alan R. Stern. Author contributions E.S. designed the MPC and performed all MPC experiments. E.S. and B.R.I. designed the MPC fabrication and performed MPC processing. E.S., A.V. and M.A.R. designed the nanosensor fabrication process and E.S., A.V. and B.R.I. performed nanosensor processing. D.J.M. assisted with MPC and nanosensor experimental design, and data analysis. E.S., A.V., N.K.R. and J.M.C. performed the sensing measurements. E.S., J.M.C. and J.P. prepared and analysed the protein samples. E.S., M.A.R. and T.M.F. wrote the manuscript and edited it, with contributions from all authors. Additional information The authors declare competing financial interests: details accompany the paper at www. nature.com/naturenanotechnology. Supplementary information accompanies this paper at www.nature.com/naturenanotechnology. Reprints and permission information is available online at http://npg.nature.com/reprintsandpermissions/. Correspondence and requests for materials should be addressed to M.A.R. and T.M.F. a 1.020 b c d 1.015 1.010 Normalized IDS Normalized IDS 1.005 1.000 0.995 1.020 2.5 ng ml−1 PSA 2.0 ng ml−1 PSA 30 U ml−1 CA15.3 15 U ml−1 CA15.3 1.04 1.03 1.02 1.01 1.00 1.015 1.010 Normalized IDS Normalized IDS 1.005 1.000 0.995 0.99 1.00 1.01 1.02 1.03 1.04 1.05 −50 0 PSA/CA15.3 spiked blood Unspiked control blood PSA PSA/CA15.3 spiked blood Unspiked control blood CA15.3 50 100 150 200 −25 25 50 75 100 125 0 Time (s) −50 0 50 100 Slope ratio = 1.38 Slope ratio = 1.94 150 200 Time (s) Time (s) −25 25 50 75 100 125 0 Time (s) Figure 4 | Label-free sensing. All sensing measurements were performed at VDS ¼ 1 V and VG ¼ –5 V and all sample introductions occurred at t ¼ 0. Normalizations were performed by dividing device currents by the pre-addition (t , 0) current level average. VDS, drain–source voltage. a, Response of an anti-PSA functionalized sensor to a MPC-purified blood sample initially containing 2.5 ng ml21 PSA (and also 30 U ml21 CA15.3), or a control sample containing neither. b, Response of an anti-CA15.3 functionalized sensor to a MPC-purified blood sample initially containing 30 U ml21 CA15.3 (and also 2.5 ng ml21 PSA), or a control sample containing neither. c,d, Normalized response of two anti-PSA (c) and two anti-CA15.3 (d) functionalized devices to MPC-purified blood containing both PSA and CA15.3, with concentrations labelled. A least-squares fit is represented by a solid black line over the selected region (line endpoints). The ratio of the normalized slopes calibrates the ratio of concentrations. LETTERS NATURE NANOTECHNOLOGY DOI: 10.1038/NNANO.2009.353 142 NATURE NANOTECHNOLOGY | VOL 5 | FEBRUARY 2010 | www.nature.com/naturenanotechnology © 2010 Macmillan Publishers Limited. All rights reserved