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Supporting information for: An electrochemical DNA-based immunoassay that employs steric hindrance to detect small molecules directly in whole blood Sahar S. Mahshid†,‡, Francesco Ricci§, Shana O. Kelley‡, Alexis Vallée-Bélisle†,* †Laboratory of Biosensors & Nanomachines, Département de Chimie, Université de Montréal, Québec H3T 1J4, Canada. §Dipartimento di Scienze e Tecnologie Chimiche, University of Rome Tor Vergata, Via della Ricerca Scientifica, Rome 00133, Italy. ‡Department of Pharmaceutical Science, Leslie Dan Faculty of Pharmacy, University of Toronto, Ontario M5S 3M2, Canada.

TABLE OF CONTENTS:

1) EXPERIMENTAL SECTION: -Mateterials, Methods; Dose-reponse curves. 2) SUPPORTING FIGURES: Figure S1. Capturing strand density validation; Figure S2. Signaling strand concentration validation; Figure S3. Optimal gain reduction; Figure S4. Dose-response curve of anti-digoxin antiserum; Figure S5. Dose-response curve of anti-TNT antibody; Figure S6. Optimization of the detection limit of the sensor; Figure S7. Detection of digoxin directly in whole blood; Figure S8. TNT CeSHHA sensor in soil.

1) EXPERIMENTAL SECTION

Materials: Polyclonal anti-Digoxin antiserum, polyclonal anti-TNT antibodies, analytical (2,4,6 Trinitrotoluene) TNT and analytical standard digoxin (Sigma-Aldrich), HPLC purified DNA strands (Biosearch Technologies Inc., Novato, CA) (Ctrl. capturing strand: 5’-HS-(CH2)6- GAGA GGG AAA AAG GAG; Sensor capturing strand (Dig): 5’-HS-(CH2)6- AAGG AAA GGG AAG AAG; Sensor capturing strand (TNT): 5’-HS-(CH2)6- TCAG TAG GAG AGT GGA; Ctrl. signaling strand: 5’- CTC CTT TTT CCC TCTC-MB; Sensor signaling strand (Dig): 5’-Digoxigenin-CTT CTT CCC TTT CCTT-MB); Sensor Signaling strand (TNT): 5’-DNP- TCC ACT CTC CTA CTGA, Tris(2-carboxyethyl) phosphine hydrochloride (TCEP; Invitrogen, Eugene, OR), 6-Mercapto-1-hexanol (MCH; Fluka) and sulfuric acid (Fisher Scientific), were all used as received. The buffer used was 50 mM NaH2PO4, 150 mM NaCl, pH 7.0. Newborn calf whole blood from Innovative research, that contains heparin as an anticoagulant, was aliquoted and frozen at -20°C prior to use.

Methods:

SCREEN PRINTED ELECTRODES were printed on transparent polyester sheets in our laboratory using a 245 DEK (Weymouth, England) screen-printing machine as previously described.[1-2] Gold-based ink (Dupont BQ331), silver ink (Electrodag 477 SS RFU) and insulating ink (Elettrodag 6018 SS) were used.

These SPEs were thoroughly rinsed in DI water, and functionalized with capturing DNA strands. 1 µL of 100 µM capturing DNA strand were incubated for 1 hr with 2 µL of 10 mM TCEP to allow the reduction of disulfide bonds and then the solution was diluted to 100 nM of capturing DNA strand in buffer media. SPEs were incubated in 40 µL of this solution on the golden working electrode area, overnight and then incubated in 3 mM MCH in buffer for another 3 hr to passivate the remaining gold area. After thoroughly rinsing with DI water, SPEs were stored in buffer. The surface density of capturing strands were determined using a previously established equation3 and estimated around 1.4 × 1012 stands/cm2.

Electrochemical measurements were performed at room temperature using a EmStatMUX potentiostat multiplexer (Palmsens Instruments, Netherland) and a standard three-electrode cell containing a platinum counter electrode (gauze-Sigma-Aldrich) and a Ag/AgCl (3M NaCl) reference electrode (CHI). The functionalized SPEs were prepared as described above using a high capturing strand density (1.4×1012 strands/cm2).[3] Experimental data were collected using square wave voltammetry from -0.05 to -0.45V in increments of 0.001V vs. Ag/AgCl, with an amplitude of 50 mV and a frequency of 60 Hz. Peak currents were fit using the manual fit mode in the PSTrace software (of Palmsens Instrument). All measurements were taken 15 min after adding antibody (digoxin antiserum or TNT antibody; 5 min

S-2 incubation) and 100nM of signaling strands (10 min incubation) to the sample containing the target molecule (digoxin or TNT). Results are presented in terms of current (knowing that the area of the electrode is 0.07 cm2). Experiments performed in whole blood were realized by adding 0.1 µL of signaling strand and 1.5 µL of anti-digoxin antiserum to 98 µL of calf whole blood. The volume percentage of the whole blood in the samples is thus 98.4%. However in case of soil, experiments were performed in aqueous soil matrix suing a 1/2 dilution of soil in buffer.

We determined the capturing strand density by employing a signaling strand that contains no ligand (Ctrl. capturing strand) and assumed that each capturing strand hybridizes to one signaling strand and that electron transfer only takes place between methylene blue and the electrode surface upon binding of the signaling DNA strand (we assumed a perfect electron transfer for each DNA bound methylene blue). Electrochemical measurements were taken following a 30-minute immersion of the sensors in a solution containing 100 nM of signaling strand to achieve saturation. The peak response corresponding to the methylene blue electron transfer (-0.25 V) was then measured and treated as reported earlier to determine the capturing strand density (knowing that the area of the electrode is 0.07 cm2).[3] Dose-response curves: Individual binding curves were fitted to a single-site binding mechanism ([T] = target concentration; Amp = current amplitude; C0 = background current; C50% is the concentration of proteins at which 50% of the sensor’s signaling amplitude was reached). LOD, limit of detection was determined as the concentration of target providing a signal that is 3 times larger than the standard deviation of the Ctrl. sample.

𝐶 " = 𝐶$ +T 𝐴𝑚𝑝T + 𝐶*$%

2) SUPPORTING FIGURES

Determination of the conditions to achieve optimal CeSHHA performance: Optimization of the CeSHHA performance was performed through a set of three calibration tests on the SPEs. We first controlled the density at which capturing DNA strands are packed on the electrode surface by varying the concentration of thiolated oligonucleotide employed during sensor fabrication. The results reveal that 100 nM (in a volume of 30 µl) is sufficient to achieve maximal surface packing, which has been shown to create maximal gain suppression upon antibody binding to the signaling strand[3] (Figure S1).

Figure S1. Optimal eSHHA performance has been observed at high capturing strand density[3]. To achieve similar performance with CeSHHA, we made sure that our SPEs were also functionalised with high capturing strand density. To do so, we incubated different SPEs overnight in a solution (30µl) containing 25 to 300 nM of capturing DNA strands. When testing these different sensors using a 100µl solution containing 100 nM of signaling strand, we found that our SPE functionalization realised with 30µl of 100 nM of capturing strand was sufficient to achieve maximal surface density. Further increasing the capturing strand density by employing for example Mg2+ typically only lead to ultra-slow signaling strand hybridization rates as reported in our previous work.[3] 100 nM of capturing strand concentration was thus employed to functionalise all our SPEs. A second experiment was then performed in order to determine the minimal concentration of signaling strand required to saturate our SPEs. To do so, we incubated our functionalised SPEs in solutions (100 µl) containing 1 to 1000 nM of signaling DNA strands. We found that 100 nM of signaling strand (100 µl) is enough to saturate most capturing strand on the sensor surface.

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Figure S2. In order to optimize the sensitivity of CeSHHA, we determined the minimal amount of signaling DNA strand that is needed to saturate all the capturing strands at the surface of the electrode. We found that 100 nM of signaling strand (in 100 µl of buffered solution) is enough to saturate most capturing strand on the sensor surface. A higher signaling strand concentration would simply reduce the detection limit of the assay by shifting the antibody binding curve (Figure S4, and the digoxin binding curve) to arbitrarily higher concentration since more proteins would be needed to saturate all signaling DNA strand in excess. Using 100 nM of signaling strands (in 100 µl of buffered solution) and a large concentration of antibodies (1/6.6 of polyclonal anti-digoxin antibody) we then confirmed that the maximal gain suppression of the sensor is achieved at high capturing strand densitity. As obserbed previously on rod electrodes,[3] we found that gain suppression indeed increases with capturing strand density and reaches a maximal value when employing a capturing strand concentration above 70 nM (which leads to a surface coverage of approximatyl 1.4×1012 strands/cm²) (Figure S3).[3]

Figure S3. Optimal gain suppression upon addition of antibody is reached at a high capturing strand density when employing capturing strand concentration above 70 nM (in 30 µl of functionalizing solution). Here we compare the signal generated by our sensor in presence of only 100 nM of signaling strand (100 µl) or in presence of 100 nM of signaling strand and 1/6.6 dilution of polyclonal anti-digoxin antibody (in 100 µl). The gain suppression was obtained using (Swith Ab – Sno Ab) / SNo Ab. The gains were measured following a 30-min acquisition.

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Figure S4. Dose-response curve of anti-digoxin antiserum, which determines the amount of antibodies required to achieve low sensor background in absence of digoxin (and thus optimal gains). 100 µl of an antiserum dilution of 1/6.6 (polyclonal anti-digoxin antibody) is sufficient to saturate the signaling strand and thus minimize the sensor background in absence of free digoxin.

Figure S5. Dose-response curve of anti-TNT antibody. 100 µl of a sample containing 350 nM of polyclonal anti-TNT antibody is sufficient to saturate the signaling strand and thus minimize the sensor background in absence of free TNT.

S-5 Figure S6. A trade-off is required in order to optimize the detection limit of the sensor: while a low concentration of antibodies renders the sensor more sensitive to the presence of low concentration of small molecule, it also increases the sensor background in absence of the analyte thus reducing the overall gain of the sensor. Here we compare the sensor’s response with different dilutions of antiserum (1/66, 1/20, 1/6.6), in presence of (A) 3 nM of target digoxin molecule (best gain -11%- obtained when using lower concentration of antibody -1/66 dilution), (B) 300 nM of target digoxin molecule (best gain -83%- obtained at higher concentration of antibodies -1/6.6 dilution).

Figure S7. Electrochemical detection of digoxin directly in whole blood. 300 nM of digoxin (green line or no digoxin: blue line), was added to newborn calf whole blood together with a 1/66 dilution of anti-digoxin antiserum. Following a 5 min incubation, 100 nM of signaling strand (black line: signaling without attached digoxingening) was added to the sample for 10 minutes before electrochemical signal acquisition. Signal gains and hybridization kinetics were found similar to the values reported in buffered solution (blue line -t1/2

no_digoxin = 20 min, green line -t1/2

+Digoxin = 19 min).

Figure S8. The TNT CeSHHA sensor can detect TNT directly in aqueous soil sample (1/2 dilution) with a C50% of 650 nM and a100-fold dynamic range.

REFERENCES

1. Ricci, F.; Amine, A.; Moscone, D.; Palleschi, G., A probe for NADH and H2O2 amperometric detection at low applied potential for oxidase and dehydrogenase based biosensor applications. Biosens Bioelectron 2007, 22 (6), 854-62. 2. Esteban Fernandez de Avila, B.; Watkins, H. M.; Pingarron, J. M.; Plaxco, K. W.; Palleschi, G.; Ricci, F., Determinants of the detection limit and specificity of surface-based biosensors. Analytical chemistry 2013, 85 (14), 6593-7. 3. Mahshid, S. S.; Camire, S.; Ricci, F.; Vallee-Belisle, A., A Highly Selective Electrochemical DNA-Based Sensor That Employs Steric Hindrance Effects to Detect Proteins Directly in Whole Blood. Journal of the American Chemical Society 2015, 137 (50), 15596-9.