Impedimetric immunosensor for the detection of circulating pro-inflammatory monocytes as infection markers

Biosensors and Bioelectronics 49 (2013) 305–311

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Impedimetric immunosensor for the detection of circulating pro-inflammatory monocytes as infection markers Armelle Montrose a,b, Sébastien Cargou c, Françoise Nepveu a,b, Rémi Manczak a,b, Anne-Marie Gué c, Karine Reybier a,b,n a

Université de Toulouse; UPS; UMR 152 Pharma-Dev; Université Toulouse 3, Faculté des Sciences Pharmaceutiques, F-31062 Toulouse cedex 09, France Institut de Recherche pour le Développement (IRD); UMR 152 Pharma-Dev; F-31062 Toulouse cedex 09, France c LAAS-CNRS, 7 avenue du Colonel Roche, F-31077 Toulouse cedex 4, France b

art ic l e i nf o

a b s t r a c t

Article history: Received 7 March 2013 Received in revised form 14 May 2013 Accepted 20 May 2013 Available online 25 May 2013

Circulating blood monocytes belong to the first line of defense against pathogens and inflammation. Monocytes can be divided into three populations defined by the expression of the cell surface molecules, CD 14 and CD 16. The CD 14++ CD 16− cells, called “classical” monocytes, represent 85% to 95% of the total monocytes in a healthy person whereas CD 14− CD 16+, called “proinflammatory” monocytes, are found in greater numbers in the blood of patients with acute inflammation and infectious diseases. This increase in the concentration of proinflammatory monocytes can be a good indicator of an infectious state. This study presents an immunosensor based on impedance detection for specific cell trapping of classical and proinflammatory monocytes. The grafting of specific antibodies (CD 14 or CD 16) was based on the use of mixed SAM associated with protein G. Each step of the functionalization was characterized by electrochemical methods, quartz crystal microbalance and atomic force microscopy. Faradaic electrochemical impedance spectroscopy and voltametric analysis confirmed the success of the modification process with a surface coverage reaching 92% for the antibody layer. The increase in the deposited mass at each step of the modification process confirmed this results revealing that one protein G in two was bound to an antibody. The cell trapping capacity, evaluated by the variation in the film resistance using non-faradaic impedance spectroscopy revealed that the cell trapping is selective, depending on the specific antibody grafted and quantitative with the range of detection being 1000 to 30,000 infected cells. This range of detection is consistent with the application targeted. & 2013 Elsevier B.V. All rights reserved.

Keywords: Immunosensor Biofunctionnalization Electrochemical impedance spectroscopy Infection diagnosis

1. Introduction Innate response against infections is related to functional activity of monocytes (MO), macrophages and polymorphonuclear leukocytes. Circulating blood MO belong to the first line of defense against infectious pathogens and inflammation (Chimma et al., 2009). Blood monocytes can be divided into three populations defined by the expression of the cell surface molecules CD 14 and CD 16: CD 14++ CD 16−, CD 14+ CD 16+, CD 14− CD 16+. The CD 14++ CD 16− cells, called “classical” MO, represent 85% to 95% of the total MO in a healthy person (Strauss-Ayali et al., 2007; Cros et al., 2010). The others, with high expression of CD 16, constitute 5% to 10% of the total MO in the blood of a healthy adult. Called proinflammatory MO, CD 16+ MO are found in larger numbers in the blood of patients with acute inflammation and infectious diseases (Auffray et al., 2009). The count of CD 14+ CD 16+ MO n Corresponding author at: Université de Toulouse; UPS; UMR 152 Pharma-Dev; Université Toulouse 3, Faculté des Sciences Pharmaceutiques, F-31062 Toulouse cedex 09, France. Tel.: +33 562259804; fax: +33 562259802. E-mail address: [email protected] (K. Reybier).

0956-5663/$ - see front matter & 2013 Elsevier B.V. All rights reserved. http://dx.doi.org/10.1016/j.bios.2013.05.025

increased from approximately 50 cells to more than 500 cells/mL in the blood of patients with severe bacterial sepsis (Fingerle et al., 1993). Consequently, a point-of-care diagnostic test detecting the level of circulating MO in blood would be helpful to the medical community for infectious diseases. The marketing of point-of-care biosensors for the detection of infectious diseases is, as explained by “the global medical device industry report 2012–2017”, the most expected application in the next ten years. Currently, the technique for the quantitative detection of CD 16+ monocytes and/or rare cells is flow cytometry. This technique has the advantage of being sensitive and reliable but is expensive, timeconsuming and not suited to both routine screening and pointof-care diagnostics. Miniaturized cell separation devices offer many advantages over conventional separation techniques (density gradient centrifugation), such as the use of small volumes, portability, low cost, improved sterile conditions in which sorting is conducted, and integration opportunities with technical analysis. Cell counting is a functional element often associated with sorting, it is generally done via optical techniques, surface plasmon resonance (SPR), and interferometric detection. Electrochemical biosensors, though label-free detection methods, constitute a promising group of sensing devices

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that afford increased sensitivities, low cost, low analysis times, affordability, and miniaturized platforms (Daniels and Pourmand, 2007). They are classified according to the observed parameter: current (amperometric), potential (potentiometric) or impedance (impedimetric). Ng et al. (2010) have developed a portable system for rapid label-free detection of endothelial progenitor cells. The system was integrated with negative dielectrophoresis for cell trapping, surface immunochemistry for selective cell capture, and impedance for detection and counting of the cells captured. Electrochemical Impedance Spectroscopy (EIS) represents a powerful method for the study of conducting materials, surface-modified electrodes and interfaces (Chang and Park, 2010). EIS uses periodic small AC amplitudes and responds to signal changes caused by the binding of target analytes to the functionalized surface of the electrodes via an immunochemical interaction. The change in the electrode impedance is measured and can be correlated to the amount of the analyte (Pejcic and De Marco, 2006). In this paper, we report the design and the characterization of an electrochemical biosensor system for quantitative detection of monocytes. The recognition between the cells and the modified electrode is based on an immunological interaction. The aim of the present study is to optimize the biosensor in view to incorporate it into a microfluidic system to develop point-of-care devices devoted to early diagnosis or therapeutic monitoring.

2. Materials and methods 2.1. Materials 2.1.1. Products 11-mercaptoundecanoic acid (MUA), 6-mercapto-1-hexanol (MH), N-hydroxysuccinimide (NHS), N-ethyl-N-(dimethylaminopropyl)-carbodiimide (EDC), phosphate buffered saline (PBS) solution, bovine serum albumin (BSA), recombinant protein G (PG), lipopolysaccharide (LPS), sulfuric acid and hydrogen peroxide were purchased from Sigma-Aldrich (Saint-Quentin Fallavier, France). Absolute ethanol was purchased from Fisher Scientific (Illkirch, France) and deionized (D.I) water was obtained using the Milli-Q water system Millipore (Molsheim, France). The CD 14 and CD 16 antibodies were purchased from Miltenyi Biotec (Paris, France). 2.1.2. Monocyte cell culture Monocytes (THP-1) were cultivated in Roswell Park Memorial Institute 1640 medium (RPMI 1640; Sigma-Aldrich) supplemented with sodium bicarbonate (1.5 g/L; Lonza (Amboise, France)), D-(+)-glucose (4.5 g/L; Sigma-Aldrich), L-glutamine (2 mM; Lonza), sodium pyruvate (1 mM; Sigma-Aldrich), 4-(2-Hydroxyethyl)1-piperazineethanesulfonic acid (10 mM; Sigma-Aldrich), 2mercaptoethanol (0.05 mM; Sigma-Aldrich) and 10% fetal bovine serum (Lonza). The cells were maintained at 37 1C in a 5% CO2 atmosphere. Monocytes were activated for 24 h by adding LPS (1 mg/mL) in the culture medium.

3. Methods 3.1. Electrode-manufacturing process Electrodes with different areas and shapes were fabricated: 1 cm2 (square), 0.5 cm2 (square), 0.04 cm2 (square) and 0.125 mm2 (rectangle). For 1 and 0.5 cm2 electrodes, a titanium/gold deposit (100 nm/ 800 nm) was layered on P-type silicium substrates (8–12 Ω/cm) by evaporation at low deposit rates (1 nm/min). Then, the substrates

were cut to obtain the electrodes with the required area. Before modification, the gold electrodes were washed in a piranha mixture (70% H2SO4/30% H2O2) for 3 min, rinsed with D.I. water and dried under a N2 stream before modification. The microelectrodes were fabricated using microelectronic massfabrication processes. The 0.04 cm2 gold electrodes were fabricated according to the process described by Torbiero et al. (2006). For the 0.125 mm2 electrodes, a titanium/platinum/gold (Ti/Pt/Au; 50/50/ 800 nm) layer was deposited on the silicium wafer. The intermediate deposit of platinum prevent the diffusion of the titanium layer into the gold and then the formation of undesirable Ti–Au alloys, while maintaining good upper interface properties (Ashwell and Heckingbottom, 1981; Martinez et al., 2010). The gold microelectrodes were passivated by a silicon nitride (Si3N4) layer of 90 nm deposited at low temperature (90 1C), to protect the electrical stripes and to define the gold sensitive area. Before use, the microelectrodes were rinsed with acetone and ethanol, rinsed with D.I. water and dried under a N2 stream. 3.2. Electrode modification The cleaned electrodes were modified as described by Ribaut et al. (2008) with few modifications notably the use of BSA to avoid nonspecific adsorption. After modification with SAM (MUA/MH 1 mM/10 mM) and then with protein G (100 g/mL in PBS) the surface was saturated by BSA for 30 min before the deposit of a commercial CD 14 or CD 16 antibody solution for 18 h at 4 1C. The electrodes were then immersed in a suspension of monocytes in PBS for 1 or 2 h depending on experiments at 37 1C in an atmosphere containing 5% CO2. To improve the probability of cell trapping on the electrodes, the incubation was carried out with stirring. 3.3. AFM imaging The measurements were acquired by AFM from Icon and were performed on dried surfaces in an acoustic mode in air at 20 1C with a humidity of 50%, using a SCANASYST-AIR triangular cantilever in peak force QNM (Quantitative NanoMechanical) tapping mode. The roughness of the topographical (Ra) was measured using the imaging processing software NanoScope Analysis from Veeco Instruments Inc. (NY, USA). 3.4. QCM analysis A quartz crystal microbalance with a 5 MHz nominal resonance frequency (Maxtek RQCM, Beaverton, USA) was used to check the modification of each layer onto a gold transducer. The quartz crystals used were covered with a 1.37 cm2 gold surface. Measurements were performed in air. 3.5. Electrochemical measurements Electrochemical experiments were performed using a Voltalab 80 PGZ 402 (Radiometer-Analytical SAS, Villeurbanne, France) and the voltamaster 4 software with a three-electrodes cell including a saturated calomel electrode (SCE) as the reference electrode, a gold electrode (1 cm²) as the counter electrode and the modified gold electrode as the working electrode. Cyclic voltammetry (CV) and faradaic electrochemical impedance spectroscopy (EIS) were performed in a PBS solution (pH¼7.4) containing 5 mM Fe2+/Fe3+ (Fe(CN)63−/4− (1:1)). In CVs, potential was cycled from −0.2 to 0.6 V with a scan rate of 100 mV/ s. The impedance spectra were recorded in a frequency range from 100 mHz to 50 kHz at the free potential of the redox couple. The amplitude of the alternating voltage was 10 mV. All experiments were performed at 25 1C in a Faraday cage.

A. Montrose et al. / Biosensors and Bioelectronics 49 (2013) 305–311

4.1.1. Electrochemical characterization Each step of the modification has been characterized by both cyclic voltammetry and faradaic impedance spectroscopy (EIS). The CV and impedimetric curves are presented in Fig. 1A and B, respectively. The Nyquist plots have been simulated according to the Randle's modified circuit (Fig. 1B). The values of the corresponding components are summarized in Table 1. The cleaned gold electrode (Fig. 1A, curve a) shows a typical CV reversible peak characteristic of the redox couple associated to a Nyquist spectrum (Fig. 1B, curve a) that follows the theoretical shape of the Randle's equivalent circuit. As expected after modification with the SAM layer, the cyclic voltammogram changes to a capacitive shape characterized by the absence of redox peak (Fig. 1A, curve b). This modification leads to the appearance of film resistance and capacitance (Rf, Qf) in the modeling (Gongadze et al., 2010) and to the increase in the charge transfer resistance. These resistance changes come from the fact that SAM form a tightly packed film that passivates the electrode and blocks the electron transfer. The increase in the double layer capacitance (Qdl) is due to the negative charges brought by the carboxylate thiolates (Xiao et al., 2011). The electrostatic repulsion of bulk ions induced by these charges is responsible for the decrease of the Warburg constant (Table 1). Replacement of the terminal carboxylate groups by amide groups suppresses the negative charges of the electrostatic repulsion blocking the electron

3.6. Statistics The statistical significance of the data has been determined using the analysis of variance (ANOVA) of the software GraphPad Prism (La Jolla, USA). 4. Results and discussion 4.1. Characterization of the modifying process The performance of impedance biosensor (selectivity, sensitivity and reproducibility) depends on the availability of recognition molecules on electrode surface. To optimize the reproducibility of the modified electrode, the antibodies have been deposited on a layer of protein G, an adhesion protein which improves the organization of the antibodies layer by specifically binding to the Fc region (Bae et al., 2005). To avoid non-specific interactions, a layer of BSA has been deposited on the protein G. The recombinant protein G lacks the albumin binding region thereby avoiding undesirable reactions with albumin. The bridge between the protein G and the gold electrode has been formed using a mixed thiol-based SAM because mixture of thiols has been shown to limit denaturation and thus to improve the bioactivity of an immobilized protein (Guiomar et al., 1999; Frederix et al., 2003; Briand et al., 2006).

5

a

4

c

900 10 Hz

800

3

- Zim / Ω cm²

Current (μA)

b

1 0 d

-1

b

700

e

f

2

307

600 10 Hz

500 f

10 Hz

400

e d

300

10 Hz

-2

c

200

0.1 Hz

-3 100 0 -5 -0.3

0.1 Hz

0.1 Hz

-4

-0.2

-0.1

0.0

0.1

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0.3

0.4

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0.7

50 kHz

0

300

0.1 Hz

0.1 Hz

a

600

900

1200 1500 1800 2100 2400 2700 3000

Zr / Ω cm²

Potential (V)

Fig. 1. (A) Cyclic voltammograms and (B) Nyquist plots of (a) a gold electrode; (b) SAM modified electrode; (c) after activation of SAM with EDC/NHS; (d) SAM/Protein G modified electrode; (e) SAM/Protein G/BSA modified electrode and (f) after immobilization of CD 14 antibody. Solution composition: 5 mM K3[Fe(CN)6]/K4[Fe(CN)6]. Scan rate 100 mV s−1. Frequency range 0.1 Hz to 50 kHz. Inset: Equivalent circuit model used for modeling the impedance spectrum characterizing the electrode/film and the film/ solution interfaces of the modified electrode.

Table 1 Values of the parameters resulting from the fitting of the Nyquist plots for each step of the modifying process, RSOL: solution resistance, Qf: film capacitance, Rf: film resistance, Rct: charge transfer resistance, Qdl: Double layer capacitance, W: Warburg constant, θ: electrode coverage. RSOL (Ω)

Qf (mF sn−1) n

Rf (Ω)

Rct (Ω)

Qdl (mF sn−1) n

W

Cleaned gold electrode

305.5 71.5

–

–

78.17 0.4

1.7197 0.008 0.636

0.0125

SAM

311.2 77.5

1.755 7 0.042 0.739

1320 7 32

939 722

6.558 7 0.15 0.924

0.0059

94.6

Activated SAM

307.4 73.4

7.270 70.08 0.909

86.5 7 0.95

6277 6.9

2.5797 0.03 0.834

0.0107

87.5

Protein G

296.6 77.1

1.829 7 0.04 0.765

316.17 7.6

7457 18

5.280 7 0.13 0.904

0.0087

89.5

BSA

308.7 74.6

1.340 7 0.02 0.742

3477 5.2

764 7 11.5

5.260 7 0.08 0.89

0.0076

89.7

295 78.8

1.7677 0.05 0.811

4337 13

914 727.4

7.936 7 0.24 0.926

0.0107

91.4

CD 14 antibody

Θ

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transfer, thus restoring the initial voltamperometric curve (Fig. 1A, curve c). For the modeling (Table 1), this phenomenon results on one hand, in an increase in the film capacitance and a decrease in the film resistance due to an increase in the film conductivity, and on the other hand, in an increase in the Warburg constant and a decrease in the charge transfer resistance due to a facilitation of the ions diffusion. Immobilization of the protein G induces a strong decrease in the current intensity (Fig. 1A, curve d) indicating that the layer blocks partially the electron transfer (“s” shaped curve). This blocking logically gives rise to an increase in the diameter of the semi-circle of the Nyquist spectrum (Fig. 1B, curve d), and thus to an increase in the charge transfer and film resistances. The increase in the double layer capacitance and the decrease in the Warburg constant can be attributed, as in the case of SAM layer, to the negative charge of the protein at pH¼7.4, as its isoelectric point (IEP) is 4.8. The decrease in the film capacitance could be attributed to the increase in the film thickness (Table 1). No significant differences were recorded after modification with BSA (Fig. 1A and B, curve e) apart from the decrease in the Warburg constant (Table 1). The binding of CD 14 antibodies (Fig. 1A, curve f) increases the peak current indicating surprisingly an increase in electron transfer. However, the Nyquist spectrum (Fig. 1B, curve f) shows an insulation of the surface with a slight increase in the charge transfer (Rtc) and film (Rf) resistances . This antagonist behavior could suggest a heterogeneous coverage of the electrode surface. An increase in the Qdl value is recorded, that may be due to the bringing of more negative charges (IEP¼ 6.1) compared to the Protein G charges whereas the film capacitance (Qf) remains of the same order. Finally, despite the presence of a more negative charged antibodies layer that should block the ion diffusion, the Warburg constant increases confirming the increase in the ions diffusion as suggested by the voltammogram. Using a redox couple as probe, impedance measurements give access to the surface coverage by evaluating the presence of pores and pinholes, through the determination of the electron transfer resistance. The surface coverage calculated from the charge transfer resistance recorded before and after modification, evolves from 95% for the SAM to 92% for the antibodies. These values are reproducible and can be used to confirm the modification of the electrode for cells detection.

4.1.2. Quartz crystal microbalance measurements The mass changes of each individual layer composing the coating have been calculated according to the Sauerbrey equation (Table 1S in Supplementary data). The corresponding number of moles absorbed per square centimetre has been calculated. The mass change recorded for the SAM layer is the most reproducible (310 ng cm−2) and significant with 2.2  10−9 mol cm−2 grafted onto the gold coated quartz crystal i.e. 1015 molecules/cm2 which is in accordance with literature (Delamarche et al., 1996). A film thickness of 1.17 nm has been calculated for the mixed SAM layer from the mass variations according to the film density (Lu and Lewis, 1972). This value is consistent with the maximum thickness of 1.6 nm described for monolayers of MUA (1 mM) adsorbed onto a gold surface (Damos et al., 2005). The activation of the terminal carboxylate groups gives rise to an increase of 589 ng cm−2. Considering the concentration of carboxylate terminated thiols in the mixed SAM layer i.e. 2.2  10−10 mol cm−2, a mass increase of 20 ng cm−2 would have been expected. The excess of EDC/NHS used to activate the carboxylate groups of the SAM and the concomitant production of secondary products such as N-acylurea (Sam et al., 2010) could be responsible for this large mass increase. Addition of the protein G produces an increase in the mass deposited of 457 ng cm−2 which corresponds to 2.1  10−11 mol cm−2 adsorbed. Finally, the variation of mass due to

Fig. 2. AFM images recorded in QNM tapping mode for each step of the modifying process: topographic image (left) and adhesion image (right). (a) Cleaned gold electrode, (b) SAM modified electrode, (c) activated SAM modified electrode, (d) protein G modified electrode, (e) BSA modified electrode and (f) antibody modified electrode.

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the immobilization of CD 14 antibodies (530 ng cm−2) and the corresponding calculated number of moles (9.6  10−12 mol cm−2) reveal that antibodies are bound to one protein G in two.

4.1.3. Microscopic observations Height topographic and adhesion images of the gold electrode obtained after each modification step are shown in Fig. 2. In order to compare the topologies of each surface, the roughness Ra has been estimated from the height images. The topographic image of the cleaned gold electrode surface shows grain boundaries with 55 nm in diameter for a roughness equal to 1.78 7 0.15 nm. The adhesion image reveals one unique stiffness with no adhesion (Fig. 2A). The surface roughness of the gold film is generally due to the gold deposition process and particularly to the nucleation generated during the crystallographic growth (Salvadori et al., 2006). The Ra of SAM layer is 1.75 70.30 nm and its surface topology is close to that of the gold surface (Delamarche et al., 1996) with grain boundaries of 60 nm in diameter (Fig. 2B). The adhesion image of SAM modified gold surface clearly reveals an increase of the adhesion which reaches 14 nN. After activation of the SAM (Fig. 2C), the Ra is 1.83 7 0.03 nm and the adhesion decreases (9 nN). The value of Ra (1.717 0.06 nm) obtained after immobilization of protein G is close to that obtained for the gold surface and for the SAM or activated SAM. Even so, the AFM images (Fig. 2D) reveal larger features and nodular structures, from 59 nm for the activated SAM layer to 80 nm for the protein G layer, and an adhesion of 4 nN. These surface modifications (activation of SAM and immobilization of protein G) depend also on the envelope profile of the gold surface and their Ra values are close to the bare gold and are coherent with the results of Bae et al. (2005). The presence of BSA (Fig. 2E) reduces the roughness (Ra ¼1.497 0.05 nm) but slightly increases the nodular structures (86 nm). The adhesion is similar to that of the Protein G layer (3.5 nN). When the antibodies are immobilized (Fig. 2F), the Ra of the dried immunosensor is 7.18 70.5 nm with features of 150 nm and the adhesion near these nodules is maximal (9 nN). All these techniques of surface characterization evidence the success of each modification step and the grafting of antibodies that will allow the cells to be trapped.

309

4.2. Characterization of the immunologic trapping 4.2.1. Detection of MO by impedance spectroscopy The capacity of the antibody modified gold electrode to trap monocytes has been first characterized using faradaic impedance spectroscopy. Incubation for 2 h of the electrode in the presence of 1 million cells per milliliter in PBS induced significant impedance variations for an electrode of 1 cm2 (curves not showed). However, the surface coverage calculated from the charge transfer resistance, deduced from modeling before and after modification, remains constant after cell trapping. This latter result suggests that more significant impedance changes would be obtained in the absence of the probe, the probe (Fe(CN)63−/4−) hindering information about cells layer. The non-faradaic (without probe) EIS curves recorded after incubation of the modified gold electrode in the PBS alone and then in presence of cells, are presented in Fig. 3 A and B, in Nyquist and Bode representations respectively. In both cases, i.e. incubation with or without cells, the Nyquist representation (Fig. 3A) highlights strong variations in impedance. In the Bode representation, a strong change in impedance is recorded at low frequencies (10 mHz to 1 Hz) after incubation in presence of cells whereas no difference appears for the incubation without cells (Fig. 3B). The change in impedance essentially due to the MO appears at frequencies lower than 100 Hz. The variation in the relative impedance Δlog(Z) (Eq. (1)) at low frequency has then been used as a valuable indicator of cells trapping efficiency. ΔlogðZÞ ¼ jlogðZðiÞÞ−logðZð0ÞÞj=logðZð0ÞÞ

ð1Þ

with Z(0) and Z(i) the impedance values deduced from the bode diagram respectively before and after monocytes trapping. The relative variation of impedance has been determined for different electrode areas (1, 0.5, 0.04 cm2 and 0.125 mm2) at 200 mHz. As expected, the miniaturization of the electrodes has resulted in a strong increase of ΔZ with values ranging from 20% for 1 cm2 to 99% for 0.125 mm2. This result is coherent with an improvement of the sensitivity for electrode size close to that of cells. However, reducing the size of the electrode could also reduce the signal-to-noise ratio and thus worsen the performances of the sensor. McAdams et al. showed that the electrodes have a noise spectrum of 1/fα with α¼ 2 n (n coefficient of the CPE), in addition to thermal noise (McAdams et al., 2006). The thermal noise, also

80000 0.1 Hz 100000

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CD 16 Antibody

log (Z)

- Zim

60000

After incubation in PBS 0.25 Hz

20000

After Monocytes trapping

CD 16 Antibody

10000

After incubation in PBS After monocytes trapping

0.1 Hz 0.5 Hz 1000 1 Hz

0

0.5 Hz

50 kHz 0

20000

40000

Zr

60000

80000

0,01

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1

10

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10000 100000 1000000

log (f)

Fig. 3. (A) Nyquist plots and (B) Bode diagrams of CD 16 antibody modified electrode, after incubation in PBS and after monocyte trapping. Solution composition: 5 mM K3[Fe(CN)6]/K4[Fe(CN)6]. Frequency range of 0.1 Hz to 50 kHz.

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1 million cells incubated. The difference between the two percentages could come from the morphological and electrical differences between healthy and activated monocytes. Indeed, CD 16+ MO are smaller and less granular than the CD 14+ ones (StraussAyali et al., 2007). Logically, this percentage strongly decreases to 33% for the trapping of healthy MO on anti-CD 16 modified electrode. This latter variation could originate either from the activation of MO caused by cells handling (undesired expression of CD 16) or from PBS ions intercalating in the film. Indeed, simple incubation of the anti-CD 14 modified electrode in PBS (2 h, 37 1C) gives rise to an impedance variation of 40% that could be ascribed to intercalation of PBS ions within the multilayer system. The statistical analysis reveals significant differences at the level of Po0.05 between the systems CD 14 and CD 16/healthy monocytes and CD 16/activated and healthy monocytes. The selectivity of the immunosensor for cells trapping is confirmed with high sensitivity towards the target cells according the antibodies immobilized and the low sensitivity towards others.

called Nyquist noise or Johnson noise, arises from the random thermal motions of charges carriers in a conducting material. The thermal noise spectrum is not dependent on the electrode surface at the thermodynamic equilibrium and is given by: DSPthermique ¼ 4KB TZ r

ð2Þ

where KB is the Boltzmann constant, T is the absolute temperature and Zr is the real impedance. To estimate the increase in noise, the noise has been calculated at 200 mHz for a cleaned gold electrode and a modified one after cell trapping for each electrode surface (Table 2S). In our case, thermal noise is low and α is similar whatever the electrode sizes, the signal-to-noise ratio remains in the same magnitude. The miniaturization allowed the increase in detection sensitivity without any increase in the noise. 4.2.2. Selectivity of trapping The selectivity of the immunosensor has been investigated by trapping healthy or activated MO on a CD 14 or CD 16 antibody modified electrode. Fig. 1S summarizes the impedance variations recorded for the different immunosensors. These results clearly demonstrate the ability of immobilized CD 14 antibody to trap monocytes and similarly for CD 16 antibody to trap infected ones. The relative impedance variation before and after cells trapping reaches 90% in the first case and 70% in the second one for

4.2.3. Analytical features of the immunosensor The impedimetric immunosensor response to various MO concentrations has also been studied. The impedance spectra of MO detection are shown in Fig. 4A in the range 1000 to 1,000,000 cells. The semicircle diameters of the Nyquist plots increase with the MO concentration. The variation of the resistance characterized by the

100000 0 1000 10000 50000 100000 1000000

80000

- Zim

60000

40000

20000

0 0

50000

100000

150000

200000

250000

Zr 350

y=145,33x - 589,84 R2 = 0.9603

300

Film Resistance

250

200

150

100

y=9,6986x + 16,447 R2 = 0.7562

50

0 0

1

2

3

4

5

6

7

-50

log[Monocytes Concentration] Fig. 4. (A) Nyquist diagrams of modified gold electrodes in PBS (pH¼ 7.4) obtained in non-faradaic mode for varying concentrations of monocytes from 0 to 106 cells/mL. The impedance spectra were recorded within a frequency range of 0.1 Hz to 50 kHz at the free potential. Electrode surface: 0.04 cm2. (B) Equivalent circuit used for modeling the interface electrode/cellular layer/solution of the modified electrode. (C) Corresponding film resistance as a function of the logarithm of the monocyte concentration.

A. Montrose et al. / Biosensors and Bioelectronics 49 (2013) 305–311

Table 2 Values of the parameters resulting from the fitting of the Nyquist plots for different cell concentrations, RSOL: solution resistance, Qf: film capacitance, Rf: film resistance, Rct: charge transfer resistance, Qdl: double layer capacitance. Concentration (cells/mL)

RSOL (Ω)

Qf (mF sn−1) n

0

448 7 76

0.517 0.09 0.903 21.057 3.6

3.717 0.6

103

372 7 11

1.23 7 0.04 0.82

36.87 1.1

3.59 7 0.11 3.85 7 0.11 0.772

104

380 7 2

6.077 0.03 0.994

45.87 0.23

5.137 0.02 3.89 7 0.02 0.785

5 104

360 7 22

1.167 0.07 0.853

76.27 4.6

1.32 7 0.08 2.03 7 0.12 0.871

105

403 7 22

1.477 0.08 0.805

1597 8.7

8.9 7 0.5

0.89 7 0.05 0.919

106

329 7 46

0.92 7 0.13 0.855 216.37 30.3 18.81 7 2.6

0.497 0.07 0.983

Rf (kΩ)

Rct (kΩ)

Qdl (mF. sn −1 )n 1.88 7 0.32 0.778

semicircle diameter, in the range of frequencies explored i.e. 100 mHz–50 kHz, appears to be a good indicator of the impedance variation according to the number of cells bound. Indeed, Cheung et al. have demonstrated that at low frequencies, the cell membrane offers a significant barrier to current flow (Cheung et al., 2010). Modeling of the Nyquist plots according to the equivalent circuit presented in Fig. 1 gave no satisfactory results. For this reason the circuit presented in Fig. 4B, generally used for metalcoated electrodes, was preferred (Ribaut et al., 2008). The values of the corresponding components are summarized in Table 2. As expected, the solution resistance is constant and the most important modification concerns the film resistance which increases with the cell concentration (Table 2) and thus with the number of cells trapped. The simulation shows also clearly two different behaviors for low and high cell concentrations. Indeed, the charge transfer resistance Rct is almost constant (∼3.9 kΩ) until 10,000 cells and then increases considerably to 18.8 kΩ for 1,000,000 cells. The film capacitance Qf also increases until 10,000 cells/mL and then decreases for higher cell concentration. These variations could be explained by the saturation of the electrode at an high cell concentration and the subsequent formation of cell multilayers. The formation of such multilayers could explain the strong decrease in the film capacity by increasing the film thickness and the strong increase in the charge transfer resistance by formation of a more compact layer (less holes). The plot of the variation of the resistance of the film as a function of the logarithm of the cell concentration highlights two distinct segments that characterize the two cell trapping modes (Fig. 4C). The intersection of the two segments at 30,000 cells is consistent with the maximum number of cells that may be attached onto an electrode of 0.04 cm2 knowing that the trapping ratio does not exceed 50%. These results demonstrate the possibility to trap and detect monocytes in the range 1000–30,000 from an homogeneous cell suspension. This detection limit is close to that described for impedimetric biosensors targeting cancer cells (Hu et al., 2013) or leukemia cells (Gu et al., 2009). Considering on one hand, the dilutions of the blood sample that will be needed before analysis in a microfluidic device and on the other hand, the trapping rate, the analytical range of the immunosensor is suitable for the detection of an increase in proinflammatory monocytes in the blood of patients. 5. Conclusion This study presents an immunosensor based on impedance detection for specific cell trapping of healthy and activated

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monocytes as a useful marker of an infectious state. The characterization of each step of the surface functionalization with specific antibodies (CD 14 or CD 16) by voltammetry, electrochemical faradaic impedance spectroscopy, quartz crystal microbalance and atomic force microscopy, has confirmed the successful antibody grafting. The cell trapping is selective since it depends on the specific antibody grafted and quantitative, with the range of detection of 1000 to 30,000 cells being consistent with the application considered. Further developments are in progress to achieve the application of point of care diagnosis from small samples of blood. Acknowledgement This work was supported by a doctoral grant from the DGA (Direction Générale de l'Armement). Appendix A. Supporting information Supplementary data associated with this article can be found in the online version at http://dx.doi.org/10.1016/j.bios.2013.05.025.

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