Optical protein sensor for detecting cancer markers in saliva
Introduction
ELISA is a standard technique to quantify the amount of protein in a solution. Protein targets are immobilized on a surface and labeled with an enzyme that can continually produce a colored product which accumulates in the solution for detection. If we track the information pathway, we find that the signal originates in the fluid sample volume, moves to the surface and returns to the volume to be finally detected. Although the signal is amplified by an enzymatic reaction, the method for taking light absorbance measurements (fluid opacity) is a relatively insensitive principle compared to optical fluorescence methods (Wells, 2006, Janasek et al., 2006).
At low target concentrations, surface immobilization (Britland et al., 1992, Mooney et al., 1996, Lahiri et al., 1999, MacBeath and Schreiber, 2000) is an efficient detection principle, especially when the surface-to-volume ratio is increased (Khandurina and Guttman, 2002, Luo et al., 2005). In this work, we examine the effects of eliminating the enzymatic signal amplification and instead, directly labeling the immobilized proteins with fluorescent probes (Espina et al., 2004, Bashir, 2004). We increase the sensitivity of detection by using fluorescence optics to measure the intensity signal (Craighead, 2006) and reducing the sample volume for increased surface-to-volume ratio. The low sample volume requirement (tens of microliters) is advantageous for clinical situations in which only a small amount of the biological fluid is available for assay.
Detecting biomarkers from a 3D volume to a 2D surface needs smaller sample volumes and can improve LOD. The challenge of this approach is the non-specific binding of non-target molecules or fluorescent probes to the surface, both of which will affect the signal integrity. Our previous development of an ultra-sensitive electrochemical RNA/DNA biosensor (Gau et al., 2001) demonstrated that a well-controlled sensor surface could reduce the background noise due to non-specific molecular binding. With this optimized surface modification protocol, detection sensitivity much less than femtomolar concentrations was achieved. This achievement effectively leads to the elimination of PCR amplification of DNA/RNA targets which is usually needed to enable detection (Wang et al., 2005). Adopting a similar surface modification approach, we have shown in this paper that IL-8 can be detected at 1.1 pM level by using a surface immobilized sandwich assay technique. In comparison, the commercial IL-8 ELISA assay has a similar LOD at 1 pM; thus, we are able to achieve the same sensitivity but without the use of enzymatic signal amplification.
IL-8 has a molecular weight of 8.5 kDa and has clinical significance for oral cancer diagnosis (St John et al., 2004). Oral cancer, the fifth most common cancer in the United States, makes up for the largest number of cancers in the head and neck category. According to the American Cancer Society, approximately 34,000 people in the United States will be diagnosed with oral cancer in 2007. The survival rate of oral cancer is 60–80% when detected during its early stages; however, this number drops to 30–40% when the cancer is diagnosed during the advanced stages (Franzmann et al., 2005). Identifying molecular markers of early disease can aid in early diagnosis, which can improve patient prognosis (Sidransky, 2002). IL-8, as a salivary biomarker for early stage oral squamous cell carcinoma (OSCC), was discovered through a previous tissue-based expression profiling effort (St John et al., 2004). Among all of the body fluids that can be tested in the clinic, saliva is the easiest to access and its collection is the least invasive to the patient. In this effort, we first use a commercial ELISA kit to execute endpoint detection of the OSCC protein biomarker from a patient's saliva; it was found that IL-8 was significantly elevated in the saliva of oral cancer patients. It was highly discriminatory in detecting oral cancer by saliva analysis. The average level of IL-8 in cancer and control patients depends on several variables, such as the number of the subjects and their physical condition. Regardless of these variables, there is an overall increase in the average measurements of IL-8 in cancer patients (Wong, 2006, Rhodus et al., 2004).
Being a filtrate of the serum, saliva has been found to have an abundance of biomarkers with major clinical significance (Herath, 2003, Sonmezoglu et al., 2005). The biomarker concentration in whole saliva can be lower than that of the serum, as in the case of TNF-α, which is another protein biomarker for oral cancer. The salivary TNF-α concentration is approximately 30 pg/ml (2 pM) in oral cancer patients and 3 pg/ml (176 fM) in healthy individuals. For detecting TNF-α in saliva, the LOD, a measure of the signal-to-noise ratio, must be several orders of magnitude lower than the baseline concentration. A need of more sensitive protein sensor is obvious. At low target concentrations, the signal is usually dominated by the background noise, which is caused by non-specific binding and optical noise. We have minimized the non-specific binding by surface modification. In optical system, the optical noise can be reduced through the utilization of confocal optics (Shotton, 1989). The use of confocal optics allows us to confine the detection volume by allowing the signal to pass through a pinhole, thereby rejecting signals that are not from the focal plane of the microscope. Indeed, a LOD as low as 4.0 fM was achieved. This value is approximately 200 times lower than the LOD for surface immobilized techniques, which lack the confocal arrangement. Our confocal optics-based surface immobilized protein sensor is essentially an ultra-sensitive protein sensor, which has potential for a wide array of applications in disease diagnostics.
In this study, three advances in optical protein biosensor are presented: (1) without enzymatic signal amplification, the surface immobilized optical protein sensor can reach a LOD of detecting IL-8 at 1.1 pM in buffer, (2) with the utilization of confocal optics, the LOD of this sensor can be further reduced by two orders of magnitude (Schweitzer et al., 2002, Zajac et al., 2007), detecting IL-8 at 4.0 fM in buffer and (3) the optical protein sensor is validated through detecting IL-8 protein in clinical saliva samples by comparing with traditional ELISA measurements. Our results show that the patients with oral cancer can be clearly distinguished from the control subjects.
Section snippets
Bio-chemicals and reagents for the optical protein micro-sensor
The same monoclonal (MAb) and polyclonal (PAb) antibodies that make up the sandwich assay pair in the human IL-8 ELISA kits were used in the detection scheme for the optical protein sensor. ELISA kits for human IL-8 protein were purchased from Pierce Endogen, Rockford, IL. The mouse anti-human IL-8 MAb, biotin-labeled (M802B), recombinant human IL-8 (RIL810), and rabbit anti-human IL-8 PAb (P801) were purchased from Pierce Biotechnology (Rockford, IL, USA). Alexa Fluor 488 labeled goat
IL-8 detection protocol by the optical protein micro-sensor
Streptavidin-coated glass cover slips were purchased from Xenopore (Hawthorne, NJ, USA). Individual sensor areas were created by adhering plastic stickers, which had pre-defined wells of approximately 60 μl capacities (Grace Bio Labs). The adhesion of the stickers to the cover slips was water-tight, preventing any cross-contamination between wells. To prime the micro-sensors for experiments, each well was incubated with 50 μl of IL-8 specific capture probe (M802B at 6 μg/ml (40 nM)) for 60 min,
Fluorescence microscopy
An epi-fluorescence microscope (Leica DMIRB) equipped with a 100 W mercury lamp and 63×, NA 0.70, dry objective was used for experiments. Fluorescence intensities from the optical protein micro-sensor surface were observed in a dark box with a Chroma filter set (Ex 480 nm, bandwidth 40 nm; Em 535 nm, bandwidth 50 nm). A 12-bit cooled monochrome CCD camera (Cool SNAP HQ, Photometrics) captured fluorescence images. The optimum exposure time was 5.0 s. ImageJ software was used for image analysis and to
Saliva sample collection and ELISA assay
Saliva samples were obtained using the same protocol as (St John et al., 2004). ELISA kits for human IL-8 (Pierce Endogen, Rockford, IL) were used according to the manufacturer's protocol. After development of the colorimetric reaction, the absorbance at 450 nm was determined by a spectrophotometer and the absorbance readings were converted to concentration based upon standard curves obtained with recombinant cytokine in each assay. Each sample was tested in at least two ELISA experiments and
T-test: statistical significance between oral cancer and healthy groups
We conducted two-tailed t-tests on the mean data of the cancer group and the control group, for both the optical protein sensor and ELISA. A two-tailed t-test (Glantz, 2002) is conducted to ensure that the difference between the two groups is not coincidental. The hypothesis that the difference is by chance, is called null hypothesis. Essentially, this test assesses whether or not the means of two groups are statistically different from each other. The groups were of moderate sample size (n = 20)
Characterization of the optical protein sensor
The limit of detection (LOD) is a figure of merit that describes the ability of a biosensor to discriminate the true signals from the noise level. In a surface-immobilized target detection scheme, each step in the protocol allows for competitive binding between the desired molecules and other non-specific molecules on the surface. Biological fluids are heterogeneous solutions and in addition to our target molecules, there are many other non-target molecules present in disproportionate numbers.
Conclusion
We have developed a platform technology that can significantly increase the sensitivity and simplify the assay preparation for protein detection. The target protein is immobilized on the surface with capture probe. The emission light from fluorophore conjugated with the reporter probe is used as the detection signal. An advantage of the presented sensor is that it can be developed with attainable materials and instrumentation for straightforward integration for point-of-care applications. For
Acknowledgment
This work is supported by the UO1 Research Grant DE15018 from the National Institute of Dental and Craniofacial Research, National Institutes of Health, Bethesda, MD 20892.
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