Biosensors and Bioelectronics 57 (2014) 59–64
Contents lists available at ScienceDirect
Biosensors and Bioelectronics journal homepage: www.elsevier.com/locate/bios
Photothermal spectral-domain optical coherence reflectometry for direct measurement of hemoglobin concentration of erythrocytes Jinyeong Yim a, Hun Kim a, Suho Ryu a, Sungwook Song b, Hyun Ok Kim b, Kyung-A Hyun a, Hyo-Il Jung a, Chulmin Joo a,n a b
School of Mechanical Engineering, Yonsei University, 50 Yonsei-ro, Seodaemun-gu, Seoul 120-749, Republic of Korea Department of Laboratory Medicine, Yonsei University College of Medicine, 50 Yonsei-ro, Seodaemun-gu, Seoul 120-749, Republic of Korea
art ic l e i nf o
a b s t r a c t
Article history: Received 30 November 2013 Received in revised form 23 January 2014 Accepted 25 January 2014 Available online 3 February 2014
A novel optical detection method for hemoglobin concentration is described. The hemoglobin molecules consisting mainly of iron generate heat upon their absorption of light energy at 532 nm, which subsequently changes the refractive index of the blood. We exploit this photothermal effect to determine the hemoglobin concentration of erythrocytes without any preprocessing of blood. Highly sensitive measurement of refractive index alteration of blood samples is enabled by a spectral-domain low coherence reflectometric sensor with subnanometer-level optical path-length sensitivity. The performance and validity of the sensor are presented by comparing the measured results against the reference data acquired from an automatic hematology analyzer. & 2014 Elsevier B.V. All rights reserved.
Keywords: Photothermal effect Hemoglobin concentration Erythrocytes Low coherence interferometry Refractive index
1. Introduction Hemoglobin (Hb) is the oxygen-transporting protein contained in the red blood cells (RBCs). The protein makes up about 97% of the RBCs, and acts as the main optical absorber and heat source in erythrocytes. The mass concentration of Hb ([Hb]) provides a measure of the oxygen-carrying capacity of blood, and can thus be used as an indicator for blood disorders and for possible complications during operation. The reference ranges of [Hb] in blood of adults are 12–16 g/dL in women and 14–18 g/dL in men (Pagana and Pagana, 2009). The thresholds of [Hb] for anemia patients are approximately 11 g/dL, 12 g/dL, and 13 g/dL for young children and pregnant women, for non-pregnant women, and for men, respectively (McLean et al., 2009). Infections with malaria cause ingestion and degradation of Hb molecules (Goldberg et al., 1990), and considerable changes in [Hb] could lead to alterations in membrane deformability, metabolic changes, hepatobiliary disease, or neurological, cardiovascular, and endocrinological disorders (Mokken et al., 1992). [Hb] monitoring is also critical to keep oxygen homeostasis in case of patients with substantial loss of blood by operation or injury, dialysis patients, women with heavy hemorrhage and menstruation, and premature births (Rosenblit et al., 1999).
n
Corresponding author. E-mail address: [email protected] (C. Joo).
http://dx.doi.org/10.1016/j.bios.2014.01.052 0956-5663 & 2014 Elsevier B.V. All rights reserved.
Because of its importance, [Hb] is measured routinely as a part of blood test. Several methods and devices for [Hb] measurement have been developed, which include the hemoglobin cyanide (cyanmethemoglobin, HiCN) method, the light-scattering technique, and the absorption difference method. The cyanide methemoglobin method employs the chemicals such as potassium cyanide (KCN) to destroy the lipid bilayer of erythrocytes. The released hemoglobin then becomes cyanide hemoglobin through cyanization and is exposed to the light of a specific wavelength for colorimetric detection (Van Kampen and Zijlstra, 1961). This method is commonly employed in the laboratories, but the toxic materials such as KCN and dimethyllaurylamine oxide are used. The light-scattering technique uses the optical properties of the red blood cell (RBC) membrane determined by a lipid bilayer. This technique, however, requires a sophisticated model of a RBC as spheroid, and therefore the measurement is highly dependent on the accuracy of the models (Mohandas et al., 1986; Tycko et al., 1985). The absorption method is not much accurate, exhibiting low correlation coefficient of o 0.804, compared with the hemoglobin cyanide method (Jeon et al., 2002). Other methods based on electrochemistry and immunoassay have also been demonstrated (John et al., 1993; Pakapongpan et al., 2011). These techniques provide highly sensitive detection of [Hb] with detection limits of o1 mg/mL. However, their detection ranges are limited to o100 μg/mL, which is not suitable for regular blood tests (Pagana and Pagana, 2009). The electrochemical techniques require sophisticated strategies for device fabrication and
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sensor surface functionalization. The immunoassay techniques have been developed to measure variants of Hbs. For instance, they have been applied to measure small amount of hemoglobin A1c, which is useful for diagnosing diabetes (John et al., 1993). Yet, commercially available assays and immunosensors require reagents and labels for operation, and involve long measurement time. Recently, Kwak et al. exploited the photothermal (PT) effect of the heme group in the RBCs. This method measured temperature changes of blood upon the light illumination to quantify [Hb] (Kwak et al., 2010a, 2010b). It provides [Hb] measurement with unprocessed blood sample, but operates with dedicated sensor chips patterned with platinum resistance temperature detectors (Pt RTDs). Here, we present a simple [Hb] detection method, based on a spectral-domain optical coherence reflectometry (SD-OCR) (Joo and de Boer, 2007). SD-OCR is based on a low-coherence interferometer, which allows identification of sensor surfaces by means of coherence gating and measures the changes of optical properties between the sensor surfaces with subnanometer optical thickness sensitivity. We employ SD-OCR to detect [Hb] in blood samples by utilizing the PT effect of Hb upon the laser illumination of a specific wavelength. The sample volume in the SD-OCR is determined by probe beam spot size and thickness of bloodcontaining chamber. We describe the implementation and performance of the photothermal SD-OCR (PT SD-OCR) sensor and demonstrate its utility as a [Hb] detector by quantifying [Hb] in the blood samples at various concentrations. Its performance is also compared with the reference measurements to assess its validity.
2. Material and methods 2.1. PT SD-OCR experimental setup A schematic diagram of PT SD-OCR sensor is depicted in Fig. 1. A 840 nm superluminescent laser diode (SLD) with a full width at half maximum (FWHM) bandwidth of 70 nm (Broadband Light Source Modules, BLM-S-840-B-I-20, Superlum, Ireland) was directed to a fiber-based Michelson interferometer. In the sample arm, the light from the fiber was collimated with a diameter of 2.4 mm and focused on a sample via an achromatic lens (AC254-035-B, Thorlabs, USA). Prior to illuminating the sample, the SD-OCR beam was combined with PT beam at a dichroic mirror (NC252113, Chroma Technology Corp., USA). A 532 nm diode pumped solid state (DPSS) laser (MGL-III-532, Continuous wave, CNI laser, China) was employed as the PT light source, because Hb molecules exhibit high
absorption at this wavelength (Lapotko and Lukianova, 2005; Lapotko et al., 2002). The absorbance of Hb at 532 nm is greater by more than two orders of magnitude than that at 840 nm. The 1/ e2 spot sizes at the sample were estimated to be 15.4 μm and 1.2 mm for SD-OCR and PT beams, respectively (Fig. 1). 2.2. Sample preparation Several control experiments were performed to evaluate the system stability. The control samples were Dulbecco's Modified Eagle's Medium (DMEM, Life Technologies, USA), Roswell Park Memorial Institute medium (RPMI 1640 Medium, Life Technologies, USA) and blood plasma separated from the whole blood with ethylenediaminetetraacetic acid (EDTA) anticoagulant. DMEM and RPMI were supplemented with 10% (v/v) fetal bovine serum (FBS, Life Technologies, USA) and 1% (v/v) penicillin streptomycin (penicillin–streptomycin, Life Technologies, USA). Experiments with blood samples were performed in compliance with the relevant laws and institutional guidelines. The samples were obtained from all voluntary donors with written informed consent. For PT SD-OCR sensor calibration, we used the blood samples provided by healthy males in the twenties. Six blood samples were acquired, and the samples were separated into erythrocytes, leukocytes and blood plasma by using a centrifuge (Fleta5, Hanil, Korea) operated at 2500 rpm for 5 min. We then prepared erythrocyte samples at 12 different Hb concentrations ([Hb]s) by diluting with autoplasma. The [Hb]s were validated with conventional hematology analyzers (Advia 2120i, Siemens AG, Germany and HemoCues Hb 201þSystem, Sweden). All the experimental sample conditions and results are shown in Table S1. A commercially available quartz cuvette (Type 49 Short part length Demountable, 49-0.1-Q, Starna Cells, UK) was used as a sample chamber. Note that any chamber made of transparent materials can be employed. The effective sample volume is determined by the SDOCR beam spot size and the height of the inner channel of the sample chamber. We filled the chamber with the sample with a volume of 15 μL, which provided sufficiently large filling for measurement. The sample volume could further be reduced down to o5 μL by use of the sample chamber with smaller volume. 2.3. Measurement method The PT SD-OCR operation is similar to that described elsewhere (Joo et al., 2005; Joo and de Boer, 2007). In brief, the sensor is based on a low-coherence spectral interferometer in which interference of reference and measurement light is spectrally
Fig. 1. A schematic of PT SD-OCR sensor for [Hb] measurement. A 840 nm superluminescent laser diode illuminates a fiber-based 2 2 coupler. The light beam in the sample arm is combined with a 532 nm photothermal (PT) beam at a dichroic mirror. The combined beams illuminate the chamber containing blood samples. The interference between the reflections from the top and bottom surfaces of inner chamber produces interference fringes. Taking the inverse Fourier transform (IFFT) of the interference spectrum results in path-length resolved complex-valued depth information, F ðzÞ. We measure the phase changes of the interference signal of our interest upon the PT beam illumination. SLD: superluminescent laser diode; DPSS: diode pumped solid state laser.
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measured, and converted into path-length resolved complexvalued information of a specimen. In the PT SD-OCR setup (Fig. 1), the SLD broadband light from the fiber illuminates the chamber, and the reflections from the surfaces of the chamber are coupled back to the fiber interferometer to produce interference spectrum at the spectrometer (USB 4000, Ocean Optics, Inc., Canada). The spectrum related to the interference can be written as pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi IðkÞ ¼ 2∑n SðkÞ Rr Rs ðzn Þ cos ð2kzn Þ; where k is the wave number and SðkÞ is the power spectral density of the light source. Rr and Rs ðzn Þ are the reflectivities of the reference surface and nth measurement surface along the beam path, respectively, with the optical path-length difference between the two surfaces zn (Joo et al., 2005). In our case, the top surface of the inner chamber was used as a reference. Taking the inverse Fourier transform of the interference spectrum with respect to 2k produces a path-length resolved complex-valued information, FðzÞ. We locate the interference signal related to the interference between the reflections from the top and bottom surfaces of inner chamber, and monitor its phase as a function of time. The phase of the interference signal at the path-length zn , ϕðzn Þ; can be evaluated as ImðFðzn ÞÞ : ϕðzn Þ ¼ tan 1 ReðFðzn ÞÞ Prior to the 532 nm PT laser illumination, no phase change is measured (Fig. 2(a)). Upon the PT laser illumination, Hb absorbs 532 nm light energy and leads to a temperature increase due to the PT effect. As the refractive index (n) varies as a function of temperature, the temperature alteration inside the chamber can be monitored by measuring the phase change of the interference signal being measured (Fig. 2(b)). The measured phase change (Δϕ) is related to refractive index variation (Δn) as
Δϕ ¼ 2k0 LΔn ¼ 2
2π
λ0
LΔn;
where k0 and λ0 are the center wave number and the wavelength of the SD-OCR light, and L is the physical distance that the light
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passes through. In our case, L corresponds to the height of the inner chamber. Measurement of phase changes due to the PT effect can be performed in either continuous or intensity-modulated illumination of PT laser. In practice, however, external disturbances such as temperature drift and vibration may affect the reliable phase measurement. Therefore, we employed the intensity-modulation scheme (Fig. 2(c)), which enables highly sensitive measurement of phase changes with improved signal-to-noise ratio (SNR). By measuring the signal around the PT beam modulation frequency only, the measurement can be done more accurately as the effect of noise components at the other frequency ranges is minimized. The intensity-modulation was achieved by using an optical chopper (Optical chopper system with the chopper wheel, MC1F2, Thorlabs, USA) in the PT beam path. For measurement, we first measured phase fluctuation for 1 min with the PT laser off as a reference. The PT laser was then turned on and modulated. The phase fluctuation was acquired at a sampling frequency of 67 Hz and the total acquisition time was 30 s. We obtained the phase fluctuation for 10 s in the middle of acquisition time. The data was then Fourier transformed, and its magnitude centered at the modulation frequency, which we here denote PT SD-OCR signal, was examined. PhaseEsensitivity depends on the SNR of the interference signal D D E 2 2 as Δϕ ¼ 1=2SNR, with the phase variance Δϕ (Joo et al., 2005). Hence it is important to obtain high SNR for sensitive measurement of PT-induced phase changes. The PT dynamics (and thus the PT SD-OCR phase signal) in time due to the absorption of the PT beam and its dependence on the modulation frequency of the laser beam have been investigated. We examined the SNR as a function of modulation frequency (Fig. S1, Supplementary information). For this measurement, a blood sample of 12.5 g/dL was used, and the irradiance of the PT laser was set to be 10.5 W/cm2. One can see that the SNR logarithmically decays with increased modulation frequency, which is consistent with earlier publications (Adler et al., 2008; Tucker-Schwartz et al., 2012). In our work, we set the modulation frequency as 1 Hz. The measurement with the modulation frequencies smaller than 1 Hz was influenced by low-frequency noise components such as
Fig. 2. [Hb] detection scheme of PT SD-OCR sensor. (a) The phase of the interference signal between the reflections from the inner surfaces of the chamber is measured in time. No change is measured prior to a 532 nm PT laser irradiation. (b) Upon the PT beam illumination, Hb molecules absorb the PT light energy and convert the photon energy into heat. The resultant temperature increase alters the refractive index of blood samples and the phase of the interference signal being measured. We quantify this phase change to measure [Hb]. (c) The phase changes under the intensity-modulated PT beam illumination. The phase signal varies at the modulation frequency. We measure the magnitude of phase signal changes at the modulation frequency only, which enables highly sensitive [Hb] measurement with improved signal to noise ratio.
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Fig. 3. (a) Representative SD-OCR phase signal with intensity-modulated PT beam illumination. (b) The magnitude of its Fourier transform (PT SD-OCR signal). The signals indicated by the stars are the higher harmonics.
Fig. 4. PT SD-OCR signals of blood plasma, leukocytes and erythrocytes with different laser intensities. The [Hb] of 15.0 g/dL was used for erythrocytes.
vibration and unexpected motion jitter of the optical chopper. Fig. 3 shows a representative phase signal, along with its corresponding PT SD-OCR signal. The sample with [Hb] of 12.5 g/dL and PT laser intensity of 9.0 W/cm2 were used.
3. Results 3.1. PT SD-OCR measurement of reference solutions Prior to performing PT SD-OCR measurement of erythrocyte samples, cell culture media were examined under the illumination of the 532 nm PT laser as the baseline. RPMI and DMEM were employed as the reference solutions, and the results are summarized in Table S1. The solutions are characterized by low absorbance at 532 nm PT laser. The magnitudes of PT SD-OCR signal of RPMI and DMEM were indeed negligible compared to the changes for erythrocytes (Fig. 4).
impossible to obtain erythrocytes-free leukocyte samples during pipetting process. For erythrocytes with [Hb] of 15.0 g/dL, PT SDOCR signals showed dramatic changes with PT laser illumination. This result indicates that Hb molecules in erythrocytes are the main contributors to PT SD-OCR signal changes. Having found that PT SD-OCR signal changes were large enough to measure [Hb], the PT SD-OCR signal changes were examined with 6 blood samples. Each blood was prepared to obtain erythrocyte samples at 12 different [Hb]s. The experimental results are shown in Fig. 5 and Table S1. It can be seen that the magnitudes of PT SD-OCR signal increase as a function of [Hb] and PT beam intensity in different manners. PT SD-OCR signals increase logarithmically with [Hb]s, whereas PT SD-OCR signals vary linearly with PT beam intensities. The logarithmic dependence of PT-SD-OCR signal on light absorber concentration has been investigated previously. For instance, Adler et al. (2008) and Tucker-Schwartz et al. (2012) performed thermal modeling and measured the logarithmic increase of photothermal phase signal on concentration of metallic nanostructures. The measured PT signals were also consistent with the results reported in Kwak et al. (2010a). We performed curve fittings for all the measurements with logarithmic and linear functions, y ¼ A lnðBx1 þ1Þ and y ¼ Cx2 . Here, y, x1 , and x2 denote PT SD-OCR signal, Hb concentration, and PT laser intensity, respectively. A–C are constants obtained from the curve fittings, and can be found in Tables S2 and S3. These results suggest that [Hb]s of blood samples can be quantitatively measured by PT SD-OCR. The detection limit of PT SD-OCR sensor was then assessed. For the [Hb] ranges in our experiment (0.4–20.0 g/dL), we first measured the noise-equivalent PT SD-OCR signals (Table S4). As the response of PT SD-OCR signal to [Hb] behaves differently with PT laser intensity (Fig. 5), the corresponding detection limits were calculated based on the calibration curves for the PT laser intensities of 7.5 W/cm2, 9.0 W/cm2 and 10.5 W/cm2. Table S4 presents the calculated [Hb] detection limits. It can be noted that [Hb] detection limit was found to be o 0.1 g/dL for the specified PT laser intensities. 3.3. Assessment of PT SD-OCR measurements
3.2. PT SD-OCR measurement of blood components We then performed the SD-OCR measurements of blood components such as blood plasma, leukocytes, and erythrocytes. As indicated in Table S1 and Fig. 4, the measurement with blood plasma exhibited no marked change. On the other hand, PT SDOCR sensor measured minute variations for the leukocyte samples. The presence of small amount of erythrocytes in the leukocyte samples was found to contribute to this small PT SD-OCR signal. The analysis with the hematology analyzer revealed that the leukocyte samples contained [Hb] of ≲0.3 g/dL. In practice, it was
In order to assess the validity of PT SD-OCR measurement and its potential applicability in clinical settings, 50 blood samples acquired from patients and healthy people were randomly chosen, and their [Hb]s were measured by PT SD-OCR sensor. The measurement were performed at different PT optical intensities (7.5 W/cm2, 9 W/cm2, and 10.5 W/cm2), and the corresponding [Hb]s, obtained based on the calibration curves described in Section 3.2, were then averaged. The [Hb]s of the blood samples were separately measured with a hematology analyzer (Advia 2120i, Siemens AG, Germany) for comparison. Fig. 6 presents the
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Fig. 5. PT SD-OCR signals as a function of [Hb] and PT laser intensity. Curve fitting constants, A–C can be found in Tables S2 and S3.
Fig. 6. Comparison of PT SD-OCR [Hb] measurements against the reference hematology analyzer (Advia 2120i).
correlation analysis between the results obtained with PT SD-OCR sensor and with the reference detector. The linear regression analysis found R2 ¼0.98, which demonstrates a great correspondence between the measurements.
4. Discussion We presented a novel optical interferometric sensor for quantifying [Hb] of blood samples. The sensor allows direct [Hb] measurement of unprocessed blood sample. The sensor demonstrated its capability in measuring [Hb] in the range of 0.4–20.0 g/ dL with a detection limit of o0.1 g/dL for the PT laser intensities of 7.5 W/cm2, 9.0 W/cm2, and 10.5 W/cm2. Noting that [Hb] of normal blood ranges from 12 to 18 g/dL (Pagana and Pagana, 2009), our sensor holds a promise in detecting the blood-related diseases such as anemia and malaria. Comparison of its performance with the hematology analyzer demonstrated a great correlation of R2 ¼0.98. Several features should be noted in our sensor. The SD-OCR sensor is based on the interference between the reflections from top and bottom surfaces of the inner chamber. The highest phase sensitivity is therefore obtained when the magnitudes of reference
and sample light are the same. As the [Hb] is higher, however, the probe beam is scattered and absorbed as the beam propagates, causing a decrease of the light reflected from the bottom surface of the inner chamber. This decrease of the reflected light intensity and the resulting imbalance between the reflections lead to a decrease in SNR of the corresponding interference signal and phase stability. This SNR decrease at high [Hb]s may account for relatively larger errors in [Hb] measurement at higher concentration range (Fig. 6). It can be resolved by using a chamber with shorter height, which reduces the path-length of light passing through the blood sample. Here, we used a sample chamber with a thickness of 100 mm, as it provided a good SNR up to [Hb] of 20.0 g/dL and a good separation of the interference signal of our interest relative to autocorrelation term. It should also be noted that PT light energy is mainly absorbed by erythrocytes in the upper part of the chamber that PT beam first passes through. As the number of erythrocytes per unit volume increases, the PT light energy is more absorbed by the cells in the upper region, and thus erythrocytes in the lower part are inhibited to absorb the light energy. Therefore, PT SD-OCR signals are expected to reach a plateau at high [Hb]s. This behavior was observed in Fig. 5 and is consistent with the earlier publication (Kwak et al., 2010a). The PT SD-OCR sensor exhibits distinct advantages and drawbacks compared with other methods. The electrochemical and immunoassay methods have demonstrated [Hb] detection with a sensitivity of o1 mg/mL, and have been employed to measure small amount of Hb variants in biological samples. However, these sensors require chemicals and dedicated sensor devices functionalized with antibodies or patterned with electrodes. Their detection ranges are also limited to o100 μg/mL. Conventional hematology analyzers enable high-precision measurement of [Hb] over the large detection range. For instance, the hematology analyzer employed in this study (Advia 2120i, Siemens AG, Germany) exhibits a precision of 0.14 g/dL and a detection range of 0–22.5 g/dL. The measurement, though, requires large sample volumes (50–200 μL), chemicals for hemolysis, and long measurement time. The photothermal sensor demonstrated by Kwak et al. (2010a, 2010b) is capable of measuring [Hb] with unprocessed blood samples without chemicals (Kwak et al. 2010a, Kwak et al. 2010b). Yet, it requires the sensor devices patterned with micro-Pt RTDs, which are costly and time-consuming to fabricate. The PT SD-OCR sensor, in contrast, does not require any functionalization of sensor surfaces and preprocessing of blood samples. The sensor chip can be any chamber made of transparent
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materials. It can operate with small sample volumes without chemicals. The smaller [Hb] is, the higher phase sensitivity can be achieved. Therefore, the sensor provides much greater sensitivity to [Hb] than absorption measurements, especially when [Hb] is small. It is characterized by the detection limit of o0.1 g/dL over the large detection range. The presented PT SD-OCR sensor was implemented with highpower light sources and a high-resolution 2048-element spectrometer. One should note, though, that the sensor can be built with less expensive, miniaturized light sources and detectors. For example, inexpensive light emitting diodes or laser diodes can be readily obtained, which provide a spectral bandwidth of more than 30 nm for SD-OCR beam and sufficient power for the PT light source. Miniature spectrometers with a smaller number of pixels are also commercially available. Implementation of PT SD-OCR sensor with these components would enable portable, chemicalsfree, and inexpensive sensing platform for [Hb] measurement. 5. Conclusions A new optical sensor for [Hb] quantification was demonstrated. The sensor utilizes a phase-sensitive optical interferometer to detect PT changes of refractive index of blood samples under the laser illumination. Our sensor is capable of measuring the [Hb]s in the blood without inducing hemolysis by chemical treatment, and is simple to implement and cost effective, as no chemicals and micro-fabrication techniques are involved. The sensor is therefore believed to be readily adopted to clinical setting, finding the relationship between [Hb]s and various blood disorders. Acknowledgments This research was supported in part by the research programs of the National Research Foundation of Korea (NRF) (MEST no. 2013-8-0622, 2013-8-1806 and 2011–0016731) and the Center for BioNano Health-Guard funded by the Ministry of Science, ICT &
Future Planning (MSIP) of Korea as Global Frontier Project (HGUARD 2013M3A6B2078959).
Appendix. Supplementary information Supplementary data associated with this article can be found in the online version at http://dx.doi.org/10.1016/j.bios.2014.01.052. References Adler, D.C., Huang, S.-W., Huber, R., Fujimoto, J.G., 2008. Opt. Express 16 (7), 4376–4393. Goldberg, D.E., Slater, A., Cerami, A., Henderson, G.B., 1990. Proc. Natl. Acad. Sci. 87 (8), 2931–2935. Jeon, K.J., Kim, S.-J., Park, K.K., Kim, J.-W., Yoon, G., 2002. J. Biomed. Opt. 7 (1), 45–50. John, W.G., Gray, M.R., Bates, D.L., Beacham, J.L., 1993. Clin. Chem. 39 (4), 663–666. Joo, C., Akkin, T., Cense, B., Park, B.H., de Boer, J.F., 2005. Opt. Lett. 30 (16), 2131–2133. Joo, C., de Boer, J.F., 2007. Opt. Lett. 32 (16), 2426–2428. Kwak, B.S., Kim, B.S., Song, S.-H., Kim, H.O., Cho, H.H., Jung, H.-I., 2010a. Analyst 135 (9), 2365–2371. Kwak, B.S., Kim, H.J., Kim, H.O., Jung, H.-I., 2010b. Biosens. Bioelectron. 26 (4), 1679–1683. Lapotko, D., Lukianova, E., 2005. Int. J. Heat Mass Transf. 48 (1), 227–234. Lapotko, D.O., Romanovskaya, T.y.R., Shnip, A., Zharov, V.P., 2002. Lasers Surg. Med. 31 (1), 53–63. McLean, E., Cogswell, M., Egli, I., Wojdyla, D., de Benoist, B., 2009. Public Health Nutr. 12 (4), 444. Mohandas, N., Kim, Y., Tycko, D., Orlik, J., Wyatt, J., Groner, W., 1986. Blood 68 (2), 506–513. Mokken, F.C., Kedaria, M., Henny, C.P., Hardeman, M., Gelb, A., 1992. Ann. Hematol. 64 (3), 113–122. Pagana, K.D., Pagana, T.J., 2009. Mosby's Manual of Diagnostic and Laboratory Tests, fourth ed. Elsevier Health Sciences, St. Louis, Missouri. Pakapongpan, S., Palangsuntikul, R., Surareungchai, W., 2011. Electrochim. Acta 56 (19), 6831–6836. Rosenblit, J., Abreu, C.R., Szterling, L.N., Kutner, J.M., Hamerschlak, N., Frutuoso, P., Paiva, T.R.S.S.d., Júnior, F., da Costa, O., 1999. Sao Paulo Med. J. 117 (3), 108–112. Tucker-Schwartz, J., Meyer, T., Patil, C., Duvall, C., Skala, M., 2012. Biomed. Opt. Express 3 (11), 2881. Tycko, D., Metz, M., Epstein, E., Grinbaum, A., 1985. Appl. Opt. 24 (9), 1355–1365. Van Kampen, E., Zijlstra, W., 1961. Clin. Chim. Acta 6 (4), 538–544.
Contents lists available at ScienceDirect
Biosensors and Bioelectronics journal homepage: www.elsevier.com/locate/bios
Photothermal spectral-domain optical coherence reflectometry for direct measurement of hemoglobin concentration of erythrocytes Jinyeong Yim a, Hun Kim a, Suho Ryu a, Sungwook Song b, Hyun Ok Kim b, Kyung-A Hyun a, Hyo-Il Jung a, Chulmin Joo a,n a b
School of Mechanical Engineering, Yonsei University, 50 Yonsei-ro, Seodaemun-gu, Seoul 120-749, Republic of Korea Department of Laboratory Medicine, Yonsei University College of Medicine, 50 Yonsei-ro, Seodaemun-gu, Seoul 120-749, Republic of Korea
art ic l e i nf o
a b s t r a c t
Article history: Received 30 November 2013 Received in revised form 23 January 2014 Accepted 25 January 2014 Available online 3 February 2014
A novel optical detection method for hemoglobin concentration is described. The hemoglobin molecules consisting mainly of iron generate heat upon their absorption of light energy at 532 nm, which subsequently changes the refractive index of the blood. We exploit this photothermal effect to determine the hemoglobin concentration of erythrocytes without any preprocessing of blood. Highly sensitive measurement of refractive index alteration of blood samples is enabled by a spectral-domain low coherence reflectometric sensor with subnanometer-level optical path-length sensitivity. The performance and validity of the sensor are presented by comparing the measured results against the reference data acquired from an automatic hematology analyzer. & 2014 Elsevier B.V. All rights reserved.
Keywords: Photothermal effect Hemoglobin concentration Erythrocytes Low coherence interferometry Refractive index
1. Introduction Hemoglobin (Hb) is the oxygen-transporting protein contained in the red blood cells (RBCs). The protein makes up about 97% of the RBCs, and acts as the main optical absorber and heat source in erythrocytes. The mass concentration of Hb ([Hb]) provides a measure of the oxygen-carrying capacity of blood, and can thus be used as an indicator for blood disorders and for possible complications during operation. The reference ranges of [Hb] in blood of adults are 12–16 g/dL in women and 14–18 g/dL in men (Pagana and Pagana, 2009). The thresholds of [Hb] for anemia patients are approximately 11 g/dL, 12 g/dL, and 13 g/dL for young children and pregnant women, for non-pregnant women, and for men, respectively (McLean et al., 2009). Infections with malaria cause ingestion and degradation of Hb molecules (Goldberg et al., 1990), and considerable changes in [Hb] could lead to alterations in membrane deformability, metabolic changes, hepatobiliary disease, or neurological, cardiovascular, and endocrinological disorders (Mokken et al., 1992). [Hb] monitoring is also critical to keep oxygen homeostasis in case of patients with substantial loss of blood by operation or injury, dialysis patients, women with heavy hemorrhage and menstruation, and premature births (Rosenblit et al., 1999).
n
Corresponding author. E-mail address: [email protected] (C. Joo).
http://dx.doi.org/10.1016/j.bios.2014.01.052 0956-5663 & 2014 Elsevier B.V. All rights reserved.
Because of its importance, [Hb] is measured routinely as a part of blood test. Several methods and devices for [Hb] measurement have been developed, which include the hemoglobin cyanide (cyanmethemoglobin, HiCN) method, the light-scattering technique, and the absorption difference method. The cyanide methemoglobin method employs the chemicals such as potassium cyanide (KCN) to destroy the lipid bilayer of erythrocytes. The released hemoglobin then becomes cyanide hemoglobin through cyanization and is exposed to the light of a specific wavelength for colorimetric detection (Van Kampen and Zijlstra, 1961). This method is commonly employed in the laboratories, but the toxic materials such as KCN and dimethyllaurylamine oxide are used. The light-scattering technique uses the optical properties of the red blood cell (RBC) membrane determined by a lipid bilayer. This technique, however, requires a sophisticated model of a RBC as spheroid, and therefore the measurement is highly dependent on the accuracy of the models (Mohandas et al., 1986; Tycko et al., 1985). The absorption method is not much accurate, exhibiting low correlation coefficient of o 0.804, compared with the hemoglobin cyanide method (Jeon et al., 2002). Other methods based on electrochemistry and immunoassay have also been demonstrated (John et al., 1993; Pakapongpan et al., 2011). These techniques provide highly sensitive detection of [Hb] with detection limits of o1 mg/mL. However, their detection ranges are limited to o100 μg/mL, which is not suitable for regular blood tests (Pagana and Pagana, 2009). The electrochemical techniques require sophisticated strategies for device fabrication and
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sensor surface functionalization. The immunoassay techniques have been developed to measure variants of Hbs. For instance, they have been applied to measure small amount of hemoglobin A1c, which is useful for diagnosing diabetes (John et al., 1993). Yet, commercially available assays and immunosensors require reagents and labels for operation, and involve long measurement time. Recently, Kwak et al. exploited the photothermal (PT) effect of the heme group in the RBCs. This method measured temperature changes of blood upon the light illumination to quantify [Hb] (Kwak et al., 2010a, 2010b). It provides [Hb] measurement with unprocessed blood sample, but operates with dedicated sensor chips patterned with platinum resistance temperature detectors (Pt RTDs). Here, we present a simple [Hb] detection method, based on a spectral-domain optical coherence reflectometry (SD-OCR) (Joo and de Boer, 2007). SD-OCR is based on a low-coherence interferometer, which allows identification of sensor surfaces by means of coherence gating and measures the changes of optical properties between the sensor surfaces with subnanometer optical thickness sensitivity. We employ SD-OCR to detect [Hb] in blood samples by utilizing the PT effect of Hb upon the laser illumination of a specific wavelength. The sample volume in the SD-OCR is determined by probe beam spot size and thickness of bloodcontaining chamber. We describe the implementation and performance of the photothermal SD-OCR (PT SD-OCR) sensor and demonstrate its utility as a [Hb] detector by quantifying [Hb] in the blood samples at various concentrations. Its performance is also compared with the reference measurements to assess its validity.
2. Material and methods 2.1. PT SD-OCR experimental setup A schematic diagram of PT SD-OCR sensor is depicted in Fig. 1. A 840 nm superluminescent laser diode (SLD) with a full width at half maximum (FWHM) bandwidth of 70 nm (Broadband Light Source Modules, BLM-S-840-B-I-20, Superlum, Ireland) was directed to a fiber-based Michelson interferometer. In the sample arm, the light from the fiber was collimated with a diameter of 2.4 mm and focused on a sample via an achromatic lens (AC254-035-B, Thorlabs, USA). Prior to illuminating the sample, the SD-OCR beam was combined with PT beam at a dichroic mirror (NC252113, Chroma Technology Corp., USA). A 532 nm diode pumped solid state (DPSS) laser (MGL-III-532, Continuous wave, CNI laser, China) was employed as the PT light source, because Hb molecules exhibit high
absorption at this wavelength (Lapotko and Lukianova, 2005; Lapotko et al., 2002). The absorbance of Hb at 532 nm is greater by more than two orders of magnitude than that at 840 nm. The 1/ e2 spot sizes at the sample were estimated to be 15.4 μm and 1.2 mm for SD-OCR and PT beams, respectively (Fig. 1). 2.2. Sample preparation Several control experiments were performed to evaluate the system stability. The control samples were Dulbecco's Modified Eagle's Medium (DMEM, Life Technologies, USA), Roswell Park Memorial Institute medium (RPMI 1640 Medium, Life Technologies, USA) and blood plasma separated from the whole blood with ethylenediaminetetraacetic acid (EDTA) anticoagulant. DMEM and RPMI were supplemented with 10% (v/v) fetal bovine serum (FBS, Life Technologies, USA) and 1% (v/v) penicillin streptomycin (penicillin–streptomycin, Life Technologies, USA). Experiments with blood samples were performed in compliance with the relevant laws and institutional guidelines. The samples were obtained from all voluntary donors with written informed consent. For PT SD-OCR sensor calibration, we used the blood samples provided by healthy males in the twenties. Six blood samples were acquired, and the samples were separated into erythrocytes, leukocytes and blood plasma by using a centrifuge (Fleta5, Hanil, Korea) operated at 2500 rpm for 5 min. We then prepared erythrocyte samples at 12 different Hb concentrations ([Hb]s) by diluting with autoplasma. The [Hb]s were validated with conventional hematology analyzers (Advia 2120i, Siemens AG, Germany and HemoCues Hb 201þSystem, Sweden). All the experimental sample conditions and results are shown in Table S1. A commercially available quartz cuvette (Type 49 Short part length Demountable, 49-0.1-Q, Starna Cells, UK) was used as a sample chamber. Note that any chamber made of transparent materials can be employed. The effective sample volume is determined by the SDOCR beam spot size and the height of the inner channel of the sample chamber. We filled the chamber with the sample with a volume of 15 μL, which provided sufficiently large filling for measurement. The sample volume could further be reduced down to o5 μL by use of the sample chamber with smaller volume. 2.3. Measurement method The PT SD-OCR operation is similar to that described elsewhere (Joo et al., 2005; Joo and de Boer, 2007). In brief, the sensor is based on a low-coherence spectral interferometer in which interference of reference and measurement light is spectrally
Fig. 1. A schematic of PT SD-OCR sensor for [Hb] measurement. A 840 nm superluminescent laser diode illuminates a fiber-based 2 2 coupler. The light beam in the sample arm is combined with a 532 nm photothermal (PT) beam at a dichroic mirror. The combined beams illuminate the chamber containing blood samples. The interference between the reflections from the top and bottom surfaces of inner chamber produces interference fringes. Taking the inverse Fourier transform (IFFT) of the interference spectrum results in path-length resolved complex-valued depth information, F ðzÞ. We measure the phase changes of the interference signal of our interest upon the PT beam illumination. SLD: superluminescent laser diode; DPSS: diode pumped solid state laser.
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measured, and converted into path-length resolved complexvalued information of a specimen. In the PT SD-OCR setup (Fig. 1), the SLD broadband light from the fiber illuminates the chamber, and the reflections from the surfaces of the chamber are coupled back to the fiber interferometer to produce interference spectrum at the spectrometer (USB 4000, Ocean Optics, Inc., Canada). The spectrum related to the interference can be written as pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi IðkÞ ¼ 2∑n SðkÞ Rr Rs ðzn Þ cos ð2kzn Þ; where k is the wave number and SðkÞ is the power spectral density of the light source. Rr and Rs ðzn Þ are the reflectivities of the reference surface and nth measurement surface along the beam path, respectively, with the optical path-length difference between the two surfaces zn (Joo et al., 2005). In our case, the top surface of the inner chamber was used as a reference. Taking the inverse Fourier transform of the interference spectrum with respect to 2k produces a path-length resolved complex-valued information, FðzÞ. We locate the interference signal related to the interference between the reflections from the top and bottom surfaces of inner chamber, and monitor its phase as a function of time. The phase of the interference signal at the path-length zn , ϕðzn Þ; can be evaluated as ImðFðzn ÞÞ : ϕðzn Þ ¼ tan 1 ReðFðzn ÞÞ Prior to the 532 nm PT laser illumination, no phase change is measured (Fig. 2(a)). Upon the PT laser illumination, Hb absorbs 532 nm light energy and leads to a temperature increase due to the PT effect. As the refractive index (n) varies as a function of temperature, the temperature alteration inside the chamber can be monitored by measuring the phase change of the interference signal being measured (Fig. 2(b)). The measured phase change (Δϕ) is related to refractive index variation (Δn) as
Δϕ ¼ 2k0 LΔn ¼ 2
2π
λ0
LΔn;
where k0 and λ0 are the center wave number and the wavelength of the SD-OCR light, and L is the physical distance that the light
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passes through. In our case, L corresponds to the height of the inner chamber. Measurement of phase changes due to the PT effect can be performed in either continuous or intensity-modulated illumination of PT laser. In practice, however, external disturbances such as temperature drift and vibration may affect the reliable phase measurement. Therefore, we employed the intensity-modulation scheme (Fig. 2(c)), which enables highly sensitive measurement of phase changes with improved signal-to-noise ratio (SNR). By measuring the signal around the PT beam modulation frequency only, the measurement can be done more accurately as the effect of noise components at the other frequency ranges is minimized. The intensity-modulation was achieved by using an optical chopper (Optical chopper system with the chopper wheel, MC1F2, Thorlabs, USA) in the PT beam path. For measurement, we first measured phase fluctuation for 1 min with the PT laser off as a reference. The PT laser was then turned on and modulated. The phase fluctuation was acquired at a sampling frequency of 67 Hz and the total acquisition time was 30 s. We obtained the phase fluctuation for 10 s in the middle of acquisition time. The data was then Fourier transformed, and its magnitude centered at the modulation frequency, which we here denote PT SD-OCR signal, was examined. PhaseEsensitivity depends on the SNR of the interference signal D D E 2 2 as Δϕ ¼ 1=2SNR, with the phase variance Δϕ (Joo et al., 2005). Hence it is important to obtain high SNR for sensitive measurement of PT-induced phase changes. The PT dynamics (and thus the PT SD-OCR phase signal) in time due to the absorption of the PT beam and its dependence on the modulation frequency of the laser beam have been investigated. We examined the SNR as a function of modulation frequency (Fig. S1, Supplementary information). For this measurement, a blood sample of 12.5 g/dL was used, and the irradiance of the PT laser was set to be 10.5 W/cm2. One can see that the SNR logarithmically decays with increased modulation frequency, which is consistent with earlier publications (Adler et al., 2008; Tucker-Schwartz et al., 2012). In our work, we set the modulation frequency as 1 Hz. The measurement with the modulation frequencies smaller than 1 Hz was influenced by low-frequency noise components such as
Fig. 2. [Hb] detection scheme of PT SD-OCR sensor. (a) The phase of the interference signal between the reflections from the inner surfaces of the chamber is measured in time. No change is measured prior to a 532 nm PT laser irradiation. (b) Upon the PT beam illumination, Hb molecules absorb the PT light energy and convert the photon energy into heat. The resultant temperature increase alters the refractive index of blood samples and the phase of the interference signal being measured. We quantify this phase change to measure [Hb]. (c) The phase changes under the intensity-modulated PT beam illumination. The phase signal varies at the modulation frequency. We measure the magnitude of phase signal changes at the modulation frequency only, which enables highly sensitive [Hb] measurement with improved signal to noise ratio.
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Fig. 3. (a) Representative SD-OCR phase signal with intensity-modulated PT beam illumination. (b) The magnitude of its Fourier transform (PT SD-OCR signal). The signals indicated by the stars are the higher harmonics.
Fig. 4. PT SD-OCR signals of blood plasma, leukocytes and erythrocytes with different laser intensities. The [Hb] of 15.0 g/dL was used for erythrocytes.
vibration and unexpected motion jitter of the optical chopper. Fig. 3 shows a representative phase signal, along with its corresponding PT SD-OCR signal. The sample with [Hb] of 12.5 g/dL and PT laser intensity of 9.0 W/cm2 were used.
3. Results 3.1. PT SD-OCR measurement of reference solutions Prior to performing PT SD-OCR measurement of erythrocyte samples, cell culture media were examined under the illumination of the 532 nm PT laser as the baseline. RPMI and DMEM were employed as the reference solutions, and the results are summarized in Table S1. The solutions are characterized by low absorbance at 532 nm PT laser. The magnitudes of PT SD-OCR signal of RPMI and DMEM were indeed negligible compared to the changes for erythrocytes (Fig. 4).
impossible to obtain erythrocytes-free leukocyte samples during pipetting process. For erythrocytes with [Hb] of 15.0 g/dL, PT SDOCR signals showed dramatic changes with PT laser illumination. This result indicates that Hb molecules in erythrocytes are the main contributors to PT SD-OCR signal changes. Having found that PT SD-OCR signal changes were large enough to measure [Hb], the PT SD-OCR signal changes were examined with 6 blood samples. Each blood was prepared to obtain erythrocyte samples at 12 different [Hb]s. The experimental results are shown in Fig. 5 and Table S1. It can be seen that the magnitudes of PT SD-OCR signal increase as a function of [Hb] and PT beam intensity in different manners. PT SD-OCR signals increase logarithmically with [Hb]s, whereas PT SD-OCR signals vary linearly with PT beam intensities. The logarithmic dependence of PT-SD-OCR signal on light absorber concentration has been investigated previously. For instance, Adler et al. (2008) and Tucker-Schwartz et al. (2012) performed thermal modeling and measured the logarithmic increase of photothermal phase signal on concentration of metallic nanostructures. The measured PT signals were also consistent with the results reported in Kwak et al. (2010a). We performed curve fittings for all the measurements with logarithmic and linear functions, y ¼ A lnðBx1 þ1Þ and y ¼ Cx2 . Here, y, x1 , and x2 denote PT SD-OCR signal, Hb concentration, and PT laser intensity, respectively. A–C are constants obtained from the curve fittings, and can be found in Tables S2 and S3. These results suggest that [Hb]s of blood samples can be quantitatively measured by PT SD-OCR. The detection limit of PT SD-OCR sensor was then assessed. For the [Hb] ranges in our experiment (0.4–20.0 g/dL), we first measured the noise-equivalent PT SD-OCR signals (Table S4). As the response of PT SD-OCR signal to [Hb] behaves differently with PT laser intensity (Fig. 5), the corresponding detection limits were calculated based on the calibration curves for the PT laser intensities of 7.5 W/cm2, 9.0 W/cm2 and 10.5 W/cm2. Table S4 presents the calculated [Hb] detection limits. It can be noted that [Hb] detection limit was found to be o 0.1 g/dL for the specified PT laser intensities. 3.3. Assessment of PT SD-OCR measurements
3.2. PT SD-OCR measurement of blood components We then performed the SD-OCR measurements of blood components such as blood plasma, leukocytes, and erythrocytes. As indicated in Table S1 and Fig. 4, the measurement with blood plasma exhibited no marked change. On the other hand, PT SDOCR sensor measured minute variations for the leukocyte samples. The presence of small amount of erythrocytes in the leukocyte samples was found to contribute to this small PT SD-OCR signal. The analysis with the hematology analyzer revealed that the leukocyte samples contained [Hb] of ≲0.3 g/dL. In practice, it was
In order to assess the validity of PT SD-OCR measurement and its potential applicability in clinical settings, 50 blood samples acquired from patients and healthy people were randomly chosen, and their [Hb]s were measured by PT SD-OCR sensor. The measurement were performed at different PT optical intensities (7.5 W/cm2, 9 W/cm2, and 10.5 W/cm2), and the corresponding [Hb]s, obtained based on the calibration curves described in Section 3.2, were then averaged. The [Hb]s of the blood samples were separately measured with a hematology analyzer (Advia 2120i, Siemens AG, Germany) for comparison. Fig. 6 presents the
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Fig. 5. PT SD-OCR signals as a function of [Hb] and PT laser intensity. Curve fitting constants, A–C can be found in Tables S2 and S3.
Fig. 6. Comparison of PT SD-OCR [Hb] measurements against the reference hematology analyzer (Advia 2120i).
correlation analysis between the results obtained with PT SD-OCR sensor and with the reference detector. The linear regression analysis found R2 ¼0.98, which demonstrates a great correspondence between the measurements.
4. Discussion We presented a novel optical interferometric sensor for quantifying [Hb] of blood samples. The sensor allows direct [Hb] measurement of unprocessed blood sample. The sensor demonstrated its capability in measuring [Hb] in the range of 0.4–20.0 g/ dL with a detection limit of o0.1 g/dL for the PT laser intensities of 7.5 W/cm2, 9.0 W/cm2, and 10.5 W/cm2. Noting that [Hb] of normal blood ranges from 12 to 18 g/dL (Pagana and Pagana, 2009), our sensor holds a promise in detecting the blood-related diseases such as anemia and malaria. Comparison of its performance with the hematology analyzer demonstrated a great correlation of R2 ¼0.98. Several features should be noted in our sensor. The SD-OCR sensor is based on the interference between the reflections from top and bottom surfaces of the inner chamber. The highest phase sensitivity is therefore obtained when the magnitudes of reference
and sample light are the same. As the [Hb] is higher, however, the probe beam is scattered and absorbed as the beam propagates, causing a decrease of the light reflected from the bottom surface of the inner chamber. This decrease of the reflected light intensity and the resulting imbalance between the reflections lead to a decrease in SNR of the corresponding interference signal and phase stability. This SNR decrease at high [Hb]s may account for relatively larger errors in [Hb] measurement at higher concentration range (Fig. 6). It can be resolved by using a chamber with shorter height, which reduces the path-length of light passing through the blood sample. Here, we used a sample chamber with a thickness of 100 mm, as it provided a good SNR up to [Hb] of 20.0 g/dL and a good separation of the interference signal of our interest relative to autocorrelation term. It should also be noted that PT light energy is mainly absorbed by erythrocytes in the upper part of the chamber that PT beam first passes through. As the number of erythrocytes per unit volume increases, the PT light energy is more absorbed by the cells in the upper region, and thus erythrocytes in the lower part are inhibited to absorb the light energy. Therefore, PT SD-OCR signals are expected to reach a plateau at high [Hb]s. This behavior was observed in Fig. 5 and is consistent with the earlier publication (Kwak et al., 2010a). The PT SD-OCR sensor exhibits distinct advantages and drawbacks compared with other methods. The electrochemical and immunoassay methods have demonstrated [Hb] detection with a sensitivity of o1 mg/mL, and have been employed to measure small amount of Hb variants in biological samples. However, these sensors require chemicals and dedicated sensor devices functionalized with antibodies or patterned with electrodes. Their detection ranges are also limited to o100 μg/mL. Conventional hematology analyzers enable high-precision measurement of [Hb] over the large detection range. For instance, the hematology analyzer employed in this study (Advia 2120i, Siemens AG, Germany) exhibits a precision of 0.14 g/dL and a detection range of 0–22.5 g/dL. The measurement, though, requires large sample volumes (50–200 μL), chemicals for hemolysis, and long measurement time. The photothermal sensor demonstrated by Kwak et al. (2010a, 2010b) is capable of measuring [Hb] with unprocessed blood samples without chemicals (Kwak et al. 2010a, Kwak et al. 2010b). Yet, it requires the sensor devices patterned with micro-Pt RTDs, which are costly and time-consuming to fabricate. The PT SD-OCR sensor, in contrast, does not require any functionalization of sensor surfaces and preprocessing of blood samples. The sensor chip can be any chamber made of transparent
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materials. It can operate with small sample volumes without chemicals. The smaller [Hb] is, the higher phase sensitivity can be achieved. Therefore, the sensor provides much greater sensitivity to [Hb] than absorption measurements, especially when [Hb] is small. It is characterized by the detection limit of o0.1 g/dL over the large detection range. The presented PT SD-OCR sensor was implemented with highpower light sources and a high-resolution 2048-element spectrometer. One should note, though, that the sensor can be built with less expensive, miniaturized light sources and detectors. For example, inexpensive light emitting diodes or laser diodes can be readily obtained, which provide a spectral bandwidth of more than 30 nm for SD-OCR beam and sufficient power for the PT light source. Miniature spectrometers with a smaller number of pixels are also commercially available. Implementation of PT SD-OCR sensor with these components would enable portable, chemicalsfree, and inexpensive sensing platform for [Hb] measurement. 5. Conclusions A new optical sensor for [Hb] quantification was demonstrated. The sensor utilizes a phase-sensitive optical interferometer to detect PT changes of refractive index of blood samples under the laser illumination. Our sensor is capable of measuring the [Hb]s in the blood without inducing hemolysis by chemical treatment, and is simple to implement and cost effective, as no chemicals and micro-fabrication techniques are involved. The sensor is therefore believed to be readily adopted to clinical setting, finding the relationship between [Hb]s and various blood disorders. Acknowledgments This research was supported in part by the research programs of the National Research Foundation of Korea (NRF) (MEST no. 2013-8-0622, 2013-8-1806 and 2011–0016731) and the Center for BioNano Health-Guard funded by the Ministry of Science, ICT &
Future Planning (MSIP) of Korea as Global Frontier Project (HGUARD 2013M3A6B2078959).
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