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. Author manuscript; available in PMC: 2016 Jun 2.
Published in final edited form as: Proc SPIE Int Soc Opt Eng. 2012 Oct 10;8460:84600S. doi: 10.1117/12.929875

Electrical detection of specific versus non-specific binding events in breast cancer cells

Benjamin C King 1, Thomas Burkhead 1, Balaji Panchapakesan 1
PMCID: PMC4890647  NIHMSID: NIHMS475181  PMID: 27274607

Abstract

Detection of circulating tumor cells (CTCs) from patient blood samples offers a desirable alternative to invasive tissue biopsies for screening of malignant carcinomas. A rigorous CTC detection method must identify CTCs from millions of other formed elements in blood and distinguish them from healthy tissue cells also present in the blood. CTCs are known to overexpress certain surface receptors, many of which aid them in invading other tissue, and these provide an avenue for their detection. We have developed carbon nanotube (CNT) thin film devices to specifically detect these receptors in intact cells. The CNT sidewalls are functionalized with antibodies specific to Epithelial Cell Adhesion Molecule (EpCAM), a marker overexpressed by breast and other carcinomas. Specific binding of EpCAM to anti-EpCAM causes a change in the local charge environment of the CNT surface which produces a characteristic electrical signal. Two cell lines are tested in the device: MCF7, a mammary adenocarcinoma line which overexpresses EpCAM, and MCF10A, a non-tumorigenic mammary epithelial line which does not. Introduction of MCF7s causes significant changes in the electrical conductance of the devices due to specific binding and associated charge environment change near the CNT sidewalls. Introduction of MCF10A displays a different profile due to purely nonspecific interactions. The profile of specific vs. nonspecific interaction signatures using carbon based devices will guide development of this diagnostic tool towards clinical sample volumes.

1. INTRODUCTION

Highly sensitive and reliable label-free biosensing is critically needed in order to advance the state of patient care, and enable the discovery of new information about cells and organisms. In current molecular biology practice, detection of molecules of interest requires tagging those molecules with an antibody or other specific recognition molecule which is labeled with a dye or other signal1. Labeling steps are often time- and reagent-consuming, and detection of the signal requires sophisticated equipment. Advances in micro and nanofabrication technology have made it feasible to construct simple devices which directly convert specific binding events to electrical or mechanical signals. Such devices can be fabricated in great quantities at low cost and could be extremely powerful for point of care diagnostics, complete patient health monitoring, and high-throughput research.

Rapid high-throughput detection is particularly in demand for purposes of cancer treatment where early detection can greatly improve survival rates and profiling the molecular signature of late-stage cancers guides physicians to select the most effective treatments2–4. Circulating Tumor Cells (CTC) are cancer cells which have entered the bloodstream5, can be accessed via a minimally invasive blood test and provide useful information on the patient’s state of health. CTCs are rare events but can be enriched by sorting procedures6–9. We are focused on the detection of cancer biomarkers in cells such as CTCs.

Single wall carbon nanotubes are ideal candidates for label-free sensing; every atom is on the surface hence their electronic properties are very sensitive to the surrounding charge environment10. The nanotubes are made sensitive to biomarkers by chemical functionalization of antibodies to the nanotube sidewalls. Many reports have described carbon nanotube sensors for cancer biomarkers as free proteins11–15, for specific DNA sequences16,17 and other molecules of interest18–20. For many of these, the sensing mechanism is the gating effect of the target proteins as they bind to the antibodies, bringing them near the CNT21,22, within the debye length of the fluid. Positively charged proteins switch off semiconducting p-type CNT, decreasing device conductance11,13–15. Negatively charged proteins switch on the CNTs, increasing the conductance12. Also, for n-type semiconducting nanowires, positively charged proteins were shown to switch them on and increase conductance13.

Cellular detection with these devices is less understood however. CNT-immunosensor detection of cells via their surface markers has only been reported with devices containing nanotubes immobilized in a 1 µm gap between electrodes. A decrease in conductance was attributed to straining of the nanotube due to the attachment of the cell23.

In order to test larger, more clinically relevant cell samples and volumes, we have developed CNT-immunosensors with a 90 µm2 area thin film. The CNT are functionalized with anti-EpCAM mouse monoclonal antibodies. Epithelial Cell Adhesion Molecule (EpCAM) is a well-studied cancer biomarker which functions in mediating homophilic cell-cell adhesions, hence its name24. EpCAM was chosen as a model system because it is present in nearly all adenocarcinomas and squamous cell carcinomas25. In the development of the sensors we observed that gas molecules in the air were doping the CNT and obscuring the sensor readings. We corrected for this and tested these sensors with EpCAM positive and EpCAM negative cells, and identified a unique electrical signature for the two different cell types. .

2. METHODS

A carbon nanotube thin film was assembled on a silicon/silicon dioxide wafer by vacuum filtration. CCVD synthesized, single wall/double wall carbon nanotube mixture (99% weight) was purchased from Cheap Tubes Inc., 1–2 nm outer diameter and 3–30 µm length. Nanotubes were suspended in IPA at 45 µg/mL and sonicated for 90 min. The solution was then diluted to 3.5 µg/mL and sonicated for 3 hours to completely disperse the nanotubes.

15 mL of the suspension was then further diluted with 85 mL of IPA and vacuum filtered over a cellulose membrane, 0.22 µm pore size. This method self-regulates the deposition rate of nanotubes on the membrane to produce an evenly distributed network26. The network was then pressed onto an oxidized (400 nm thickness) silicon wafer for 30 minutes. Next the wafer was transferred to an acetone vapor bath to dissolve the membrane.

The targeted thin film density was <5 CNT/µm, while still yielding stable and reproducible devices. Very thin films leave more nanotube sidewalls exposed and reduce the number of junctions per nanotube so that nanotube resistance is more significant relative to the junction resistance27,28. Both of these attributes improve the sensing properties of the film. SEM images were surveyed to determine the film density (Figure 2.1). Mean density of 3.93 CNT/µm was found by counting nanotube intersections with a grid of 1 µm lines superimposed on the SEM image. The distribution of the nanotubes was favorable as well. The standard deviation was 3.22, therefore over 68% of the film area is favorable for sensing, containing 1–8 CNT/µm.

Figure 2.1.

Figure 2.1

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Left: SEM image of the CNT thin flim showing well dispersed nanotubes. Right: Histogram of CNT/µm counted in a 20×15 µm area

Patterning of the nanotube film and electrode and insulating layer fabrication were done by photolithography in the cleanroom. AZ4620 photoresist is used to mask the nanotube film areas needed for the sensor elements. Exposed nanotubes are etched away in a March reactive ion etcher for 90 s at 200 W power and 10% O2. SC1827 photoresist is used to mask the electrode pattern. Electrodes consist of a 10 nm Ni adhesion layer and a 90 nm Au layer. They are deposited by sputtering (Leskar PVD 75 system, 300 W DC power). Lastly, the sensors are covered with SU8–2005, a 5 µm thick photopolymer layer. A window over each of the nanotube sensor elements is developed but the electrodes remain insulated beneath the SU8.

Finished carbon nanotube sensors were functionalized with anti-EpCAM by a pyrene linker molecule. The pyrene rings of 1-Pyrenebutanoic acid, succinimidyl ester (PASE) adsorb onto carbon nanotube sidewalls by π-stacking29. The ester on the other end of the molecule provides an attachment point for antibodies. PASE (AnaSpec, #81238) was dissolved in methanol at 1 mg/mL. Devices were incubated in the PASE solution for 1 hour at room temperature, then rinsed with methanol and water. Devices were then incubated in anti-EpCAM (10 µg/mL in PBS, EMD Bioscience, #OP187) for 1 hour at 37°C. Lastly, a surfactant, Tween20, was used to block unfunctionalized nanotube sidewalls or PASE sites in order to minimize nonspecific interactions. Devices were incubated in 0.5% Tween20 for 2 hours at room temperature. After incubation, devices were washed with water, then each incubated in a 2 µL droplet of PBS overnight in a humid chamber before testing.

Tests were performed on MCF-7 and MCF-10A cell lines, purchased from American Type Culture Collection. MCF-7 is a breast adenocarcinoma cell line which is EpCAM positive and MCF-10A is a non-tumorigenic breast epithelial cell line which does not express EpCAM. Confocal microscopy was used to verify the EpCAM expression of the MCF-7 cells (Figure 2.3).

Figure 2.3.

Figure 2.3

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MCF-7 (left) stained for nucleus (blue) and EpCAM (green). MCF-10A (right) with the same staining shows negligible EpCAM signal

Devices were placed on a Signatone probe station and the probe tips contacted the device at the source and drain terminals. These were wired to an Agilent 4156C Semiconductor Parameter Analyzer which was controlled via a custom LabVIEW interface, running on a windows PC. A 100 mV bias was applied and the source drain current, ISD, was recorded for the duration of each test. ISD for the devices as fabricated typically ranged from 250 to 400 nA. To compare results among devices, ISD data were normalized to the initial value obtain the G/G0 signal. In cases where the current drifted before addition of the sample, signals were normalized to the steady state value before sample addition (denoted as G/GSS).

Test solutions and samples were pipetted directly onto the device. In some cases the test was initiated with the device already hydrated and samples were pipetted into this droplet. The final testing protocol started with the device hydrated in a 2 µL droplet which was placed immediately after functionalization and left overnight. The bias was applied, and then a 5 µL droplet, either of cell suspension, PBS or water, was pipetted directly into the standing 2 µL droplet. The sensor element was also viewed an optical microscope to confirm the presence of cells. A concentration of 10,000 cells/µL consistently covered the sensor element with 20–30 cells (Figure 2.2). Multiple MCF-7 and MCF-10A samples were tested at each session, along with blank PBS and DI water samples.

Figure 2.2.

Figure 2.2

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Top Left: Sensor schematic, nanotube film on wafer with electrodes and insulating SU8 layer. Top right: MCF10A cells on the surface of the sensor during testing. Bottom: Biofunctionalization schematic of the nanotubes

3. RESULTS

3.1. Sensor Development

In the initial test protocol, the sensor, dry and exposed to air, was probed and a 100 mV bias applied for 300 seconds. At 30 seconds, a sample of cells suspended in PBS was pipetted onto the device. Wetting the sensor with the buffer usually attenuated the current as expected. Unfortunately, this attenuation was very inconsistent from device to device and about one in ten devices produced an increase or random fluctuations in conductance. Thus any signals arising due to specific binding were masked by the sudden and drastic change in the sensor environment. Many repetitions failed to generate a discernible pattern, or a characteristic difference between the specific and nonspecific binding signatures (Figure 3.1).

Figure 3.1.

Figure 3.1

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Effects of liquid addition and incubation on sensors. Addition of cells suspended in PBS to dry sensors (left) caused inconsistent conductance decreases, regardless of cell type. Attempts to observe specific interactions (right) over extended times also failed to show a significant difference between the two cell types, but did suggest that extended exposure to the solutions caused the conductance to decrease.

Placing a PBS droplet on the device before or after applying the VDS bias removed the initial-hydration current attenuation from the signal. However, noisy and anomalous readings persisted upon addition of the cell samples. Reproducible signals remained impossible to obtain.

Attempting to avoid the noise associated with the liquid sample introduction, a study was conducted over an extended time interval. Cell samples at a concentration of 10,000 cells/µL were added to the devices and placed incubated at 37°C. A current measurement was taken at 2 hours, 4 hours and 8 hours (Figure 3.1). Then the cells were fixed in methanol, incubated in PBS overnight, and another reading was taken the next day. There was no significant difference between the MCF7 and MCF-10A samples. But the pattern which emerged saw the current in all devices increase slightly up to the 4 hour time point, then decrease at the 8 hour and post fixation time points to below the initial value. This indicated that incubation in aqueous solutions alone was altering the nanotube conductance.

Tests on bare, non-functionalized carbon nanotube devices revealed that extended incubation in water decreased device conductance (Figure 3.2). Furthermore, device conductance gradually increased as they were exposed to air following fabrication. This effect was reversed by storing the fabricated sensors in a vacuum chamber.

Figure 3.2.

Figure 3.2

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Environmental effects on non-functionalized nanotube sensors. Left: Addition of DI water to the blank nanotube film decreases its conductance in subsequent tests. Right: Average thin film conductance increases by 12% after 2 hours of air exposure

Altogether these findings indicate that gas molecules were reversibly doping the carbon nanotubes as reported previously30. The effect accumulated as they were exposed to air, and was reversed for vacuum and liquid environments. When in these environments, high concentrations of gas molecules in the vicinity of the nanotubes would either be drawn out into the vacuum or dissolved into the liquid. The reversing of the doping effects is not consistent among sensors and lead to erratic behavior.

Therefore, a fabrication and functionalization protocol was developed which minimizes the nanotubes exposure to air as soon as the cleanroom fabrication steps are complete. The CNT film was repeatedly coated with photoresists and washed in developer solution during fabrication so it is assumed that t=0 for air exposure at the moment the last photolithography step is complete.

Devices constructed with this minimal air exposure method produced much more consistent responses to blank PBS solutions for one hour (Figure 3.3). Additional PBS droplets had to be added to prevent the devices from drying out during the test. Bare and functionalized devices both exhibited similar curve shapes, a sharp drop in current followed by gradual recover. However the functionalized device current gradually declined over the interval unlike the bare devices.

Figure 3.3.

Figure 3.3

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Devices prepared with minimal air exposure generate reproducible signals upon addition of PBS, with and without antibody functionalization

Consistent performing devices provided a stable platform to investigate the specific binding signature, although one small issue remains unresolved. When the bias is applied, some device currents drift upwards, some downwards, and some are steady. All three examples are shown in Figure 3.3 on the right. It is unclear what effect, if any, this has on device sensitivity.

3.2. Specific Signature Identification

Elimination of noise due to gaseous doping enabled clear identification of the specific vs. nonspecific binding signatures by testing 9 devices, 4 MCF-7, 4 MCF-10A, and 1 PBS control. The MCF-10A signals all had the characteristic shape for plain PBS addition as established previously. The PBS control reproduced this shape as well. For the MCF-7 samples, 3 of the 4 exhibited an inverse shape while 1 had the nonspecific shape.

To determine statistical significance, the average slope of the 0.4 seconds immediately following the prominent inflection point was calculated for each signal. Both data sets were determined to be random and of equal variance, therefore a two-sample t-test was used. The p value of 0.024 < 0.05 denotes that there is a significant difference between the sampled MCF-7 signals and MCF-10A signals.

Both cell types were also tested on blank films (Figure 3.4), and typically displayed the characteristic nonspecific shape except for one outlier. The same statistical calculation found that the two data sets were not significantly different. This shows that the sensor is capable of differentiating between two cell populations which are very similar except for their surface markers, the first such accomplishment for a carbon nanotube-thin film electrical sensor.

Figure 3.4.

Figure 3.4

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Characteristic signals. Sample addition causes a temporary response in the sensor conductance. For sensors functionalized with anti-EpCAM (left), the EpCAM negative control cells cause a sudden decrease in conductance and gradual return to approximately the previous value. The same signal morphology is observed for addition of pure PBS to the device. EpCAM positive cells typically generate the inverse signal shape. For non-functionalized (blank) CNT surfaces, both cell types typically produced local minima in the conductance signals. The readings from the blank sensors were similar to the negative controls for the functionalized sensors. All curves are normalized to their steady state conductance and are plotted spaced apart for clarity.

4. DISCUSSION

The MCF7 and MCF10A cell lines, like most mammalian non-excitable cells have a negative resting potential. The potential is maintained by active transport of ions, mainly sodium and potassium, across the membrane in order to facilitate transport of other chemical species. MCF7 potential was reported to vary from -58.6 mV to -2.7 mV with the cell cycle31. MCF10A membrane potential is estimated to be about −10 mV32 but a similar report correlated to the cell cycle was not available.

As MCF7 cells reached the sensor and EpCAM units bound to the anti-EpCAM of the sensor, the sharp increase in conductance suggests the negatively charged MCF7s influenced the local charge environment to switch on the nanotubes in a manner similar to that observed for negatively charged free protein12. The sensor current then began to gradually return to near its initial value, either due to a loss of charge by the cell or leakage through the antibody-PASE linkage binding the cell to the antibodies. We did not observe a conductance decrease as reported before for very small channel sensors23 because the CNT networked in a thin film are free to move about and would not experience such a strain due to antibody-antigen binding.

Introduction of MCF10A to the device produces an opposite shaped curve, of the same shape as for blank PBS addition, supporting the hypothesis that specific interaction is required for the cell to affect the device. The cause of the shape of the curve produced by the nonspecific MCF10A samples and blank PBS with no interaction is uncertain. Perhaps it is due to the agitation of the fluid upon pipetting.

Controls with the non-functionalized films did not show a significant difference between the MCF-7 and MCF-10A cell types so the different signals for the functionalized sensors cannot be due solely to a difference in membrane potential. However, a thorough comparison of the membrane potentials of the MCF7 and MCF10A lines would more strongly confirm these findings. Tumorigenesis and cell cycle deregulation are often closely linked33,34. If the MCF-7 cell cycle is sufficiently different from the MCF-10A, it could result in the two cell populations having different membrane potential distributions. The cell cycle dependent membrane potential of the MCF7 may also cause the signal to vary among tests based on the net membrane potential of the 20–30 cells interfacing with the device for a given experiment.

5. CONCLUSIONS

This study shows, for the first time, sensing of cellular EpCAM by a thin film carbon nanotube device. The sensing is performed entirely on a simple two terminal device with no moving parts. Such simplicity would enable this type of device to readily be integrated with other micro/nano components for biotechnology. It is possible to envision these types of sensors in many applications: Remote areas or those without high-tech infrastructure, home use for patients who need to regularly monitor CTC counts or other levels, and even perhaps in implantable devices, wirelessly relaying real time data to physicians and care providers.

Development of the sensor throughout this study encountered and overcame many pitfalls, most notably the gas adsorption and desorption effects. The steps needed to control these errors were very simple. Minimizing device exposure to air and keeping the device constantly hydrated from the end of the biofunctionalization process to the initiation of testing transformed erratic unpredictable sensors into reliably performing ones.

Some characteristics of the electrical signals have yet to be explained and are under continued investigation. Fully reproducible carbon nanotube biosensors remain an important goal35. Future work will also repeat tests as described here for other cancer markers and cell types, pushing towards development of multiplexed CNT biosensors which could give complete profiles of cancer markers and more.

Table 3.1.

Statistical comparison between cell type signals for the functionalized sensor

MCF-7 MCF10-A
Calculated slopes after sample addition −0.1469 0.162
−0.0668 0.238
−0.0601 0.0726
0.0395 0.0291
Average −0.0586 0.1254
Std. Deviation 0.0764 0.0932
t statistic −3.05326
Critical +/−2.47015
p value 0.02352
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Acknowledgements

The authors thank Thomas Rousell for developing the LabVIEW testing interface and Vanessa Velasco for design of the device photomasks. Funding for this research was provided by National Cancer Institute of the National Institute of Health, Grant Number: 1R15CA156322-01A1 for Balaji Panchapakesan.

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