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Unconstrained pulse pressure monitoring for health management using hetero-core fiber optic sensor MICHIKO NISHIYAMA,* MASAKO SONOBE, AND KAZUHIRO WATANABE Department of Science and Engineering for Sustainable Innovation, Faculty of Science and Engineering, Soka university, 1-236 Tangi-Machi, Hachioji, Tokyo, 192-8577, Japan * [email protected]
Abstract: In this paper, we present a pulse pressure waveform sensor that does not constrain a wearer’s daily activity; the sensor uses hetero-core fiber optics. Hetero-core fiber sensors have been found to be sensitive to moderate bending. To detect minute pulse pressure changes from the radial artery at the wrist, we devised a fiber sensor arrangement using three-point bending supports. We analyzed and evaluated the measurement validity using wavelet transformation, which is well-suited for biological signal processing. It was confirmed that the detected pulse waveform had a fundamental mode frequency of around 1.25 Hz over the timevarying waveform. A band-pass filter with a range of frequencies from 0.85 to 1.7 Hz was used to pick up the fundamental mode. In addition, a high-pass filter with 0.85 Hz frequency eliminated arm motion artifacts; consequently, we achieved high signal-to-noise ratio. For unrestricted daily health management, it is desirable that pulse pressure monitoring can be achieved by simply placing a device on the hand without the sensor being noticed. Two types of arrangements were developed and demonstrated in which the pulse sensors were either embedded in a base, such as an armrest, or in a wearable device. A wearable device without cuff pressure using a sensitivity-enhanced fiber sensor was successfully achieved with a sensitivity of 0.07–0.3 dB with a noise floor lower than 0.01 dB or multiple subjects. ©2016 Optical Society of America OCIS codes: (060.2370) Fiber optics sensors; (170.3890) Medical optics instrumentation.
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T. Tamura, T. Togawa, M. Ogawa, and M. Yoda, “Fully automated health monitoring system in the home,” Med. Eng. Phys. 20(8), 573–579 (1998). M. Engin, A. Demirel, E. Z. Engin, and M. Fedakar, “Recent developments and trends in biomedical sensors,” Measurement 37(2), 173–188 (2005). M. Folke, L. Cernerud, M. Ekström, and B. Hök, “Critical review of non-invasive respiratory monitoring in medical care,” Med. Biol. Eng. Comput. 41(4), 377–383 (2003). X. Zhu, W. Chen, T. Nemoto, Y. Kanemitsu, K. Kitamura, K. Yamakoshi, and D. Wei, “Real-time monitoring of respiration rhythm and pulse rate during sleep,” IEEE Trans. Biomed. Eng. 53(12 Pt 1), 2553–2563 (2006). W. Chen, X. Zhu, T. Nemoto, Y. Kanemitsu, K. Kitamura, and K. Yamakoshi, “Unconstrained detection of respiration rhythm and pulse rate with one under-pillow sensor during sleep,” Med. Biol. Eng. Comput. 43(2), 306–312 (2005). S. Ogoh, “Middle cerebral artery flow velocity and pulse pressure during dynamic exercise in humans,” Heart Circul. Physiol. 288(4), H1526–H1531 (2004). R. A. Payne, C. N. Symeonides, D. J. Webb, and S. R. J. Maxwell, “Pulse transit time measured from the ECG: an unreliable marker of beat-to-beat blood pressure,” J. Appl. Physiol. 100(1), 136–141 (2006). M. Nitzan, S. Turivnenko, A. Milston, A. Babchenko, and Y. Mahler, “Low-frequency variability in the blood volume and in the blood volume pulse measured by photoplethysmography,” J. Biomed. Opt. 1(2), 223–229 (1996). S. Tanaka, Y. Matsumoto, and K. Wakimoto, “Unconstrained and non-invasive measurement of heart-beat and respiration periods using a phonocardiographic sensor,” Med. Biol. Eng. Comput. 40(2), 246–252 (2002). M. Rothmaier, B. Selm, S. Spichtig, D. Haensse, and M. Wolf, “Photonic textiles for pulse oximetry,” Opt. Express 16(17), 12973–12986 (2008). T. Allsop, K. Carroll, G. Lloyd, D. J. Webb, M. Miller, and I. Bennion, “Application of long-period-grating sensors to respiratory plethysmography,” J. Biomed. Opt. 12(6), 064003 (2007). A. Babchenko, B. Khanokh, Y. Shomer, and M. Nitzan, “Fiber optic sensor for the measurement of respiratory chest circumference changes,” J. Biomed. Opt. 4(2), 224–229 (1999). S. Sprager and D. Zazula, “Heartbeat and respiration detection from optical interferometric signals by using a multimethod approach,” IEEE Trans. Biomed. Eng. 59(10), 2922–2929 (2012).
#266540 Journal © 2016
http://dx.doi.org/10.1364/BOE.7.003675 Received 18 May 2016; revised 14 Aug 2016; accepted 22 Aug 2016; published 26 Aug 2016
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14. D. Lau, Z. Chen, J. T. Teo, S. H. Ng, H. Rumpel, Y. Lian, H. Yang, and P. L. Kei, “Intensity-modulated microbend fiber optic sensor for respiratory monitoring and gating during MRI,” IEEE Trans. Biomed. Eng. 60(9), 2655–2662 (2013). 15. L. Dziuda, F. W. Skibniewski, M. Krej, and J. Lewandowski, “Monitoring Respiration and Cardiac Activity Using Fiber Bragg Grating-Based Sensor,” IEEE Trans. Biomed. Eng. 59(7), 1934–1942 (2012). 16. M. Nishiyama, H. Sasaki, and K. Watanabe, “A deformation sensitive pad-structure embedded with hetero-core optic fiber sensors,” Sens. Actuators A Phys. 136(1), 205–211 (2007). 17. M. Nishiyama and K. Watanabe, “Wearable sensing glove with embedded hetero-core fiber-optic nerves for unconstrained hand motion capture,” IEEE Trans. Instrum. Meas. 58(12), 3995–4000 (2009). 18. M. Nishyama, M. Miyamoto, and K. Watanabe, “Respiration and body movement analysis during sleep in bed using hetero-core fiber optic pressure sensors without constraint to human activity,” J. Biomed. Opt. 16(1), 017002 (2011). 19. H. Sasaki, Y. Kubota, and K. Watanabe, “Sensitivity property of a hetero-core splice fiber optic displacement sensor,” Proc. SPIE 5579, 136–143 (2004). 20. J. K.-J. Li, Dynamics of the Vascular System (World scientific publishing, Singapore, 2004).
1. Introduction In recent years, the medical welfare network has advanced to an extent where diseases can be prevented before their occurrence [1–3]. Therefore, health monitoring techniques that people can perform themselves on a continuous basis are preferred, particularly, pulse pressure monitoring implemented for everyday activities such as sleeping and exercising [4–6]. These health monitoring devices should allow for continuous and autonomous detection of vital physiological indicators while being portable and comfortable from a user’s viewpoint. In other words, the user should not notice that they are wearing a sensor. In addition, the system must be robust to environmental disturbances such as electromagnetic interference (EMI), external temperature, and pressure fluctuations. Currently, there are several conventional methods for measuring someone’s pulse and blood pressure. However, an electrocardiogram (ECG) [7] or a photoelectric plethysmogram [8] may bring discomfort and inconvenience to the subject and physician because the sensors must be placed on the body. In addition, these techniques also require appropriate sensors to be installed on the subject. Conventional noninvasive blood flow pressure measurement is based on an inflatable pressure cuff surrounding an arm or a finger. However, periodic interruption of blood flow can affect a user’s physiological state and prevent comfortable sleep because this method requires inflation and deflation of the cuff. To meet the requirements for pulse measurement to be noninvasive and unconstrained, a low-cost pillow-shaped respiratory monitor [5] and a phonocardiographic sensor set on an air mat or a water mat that subjects sleep on [9] have been proposed. However, the raw signals generated using these unconstrained methods were noisy because of the inherent vibrational signals in the sensor output; consequently, an algorithm based on a wavelet transformation or appropriate band-pass filter was needed to be adapted to the raw signal. In the past few decades, several techniques have been developed for monitoring human respiration using optical fiber sensors, such as motion capture and monitoring systems [10–15]. This is because optical fiber sensors have several advantages, such as the sensor not needing a power supply, resistance to EMI, flexibility, light weight, and high resistance to moisture. In one example it was proposed to use plastic optical fibers for pulse oximetry [10]. However, plastic optical fibers have limited multi-mode stability for bending actions in the transmission line. Other approaches for respiration and heartbeat monitoring were employed, such as using optical interferometers with optical fiber length changes [13], microbend fiber sensors based on intensity modulation [14], and fiber Bragg gratings [15]. Although these techniques utilized an optical fiber, the optical interferometer methods were easily affected by external disturbances along the transmission fiber line because of the focus on the fiber length change owing to external perturbations. The microbend technique may induce overstrain in the fiber, and consequently, the initial transmission loss may seem to increase and the signal-to-noise ratio may decrease. Fiber Bragg gratings are well-known as strain and temperature sensors; however, they inherently need to compensate for the environmental temperature fluctuations for strain sensing. We have developed hetero-core optical fiber sensors [16–19] that can detect moderate bending with high sensitivity to the optical light intensity. The use of hetero-core optical fiber sensors
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has been proposed for displacement and pressure sensing in environmental monitoring, as well as for wearable motion analysis [17, 18]. This is because a hetero-core optical fiber sensor is sensitive only to the bending of the sensor portion. The fiber transmission line is unaffected by external disturbances such as pressure caused by the single-mode (SM) stable propagation scheme. Additionally, the fiber transmission line is independent of temperature fluctuations inside the sensor. Therefore, hetero-core optical fiber sensors have the potential to be used for unconstrained pulse pressure monitoring with high signal-to-noise ratio because of their stability to external disturbances. Such a health monitoring system can be simply constructed without the complex algorithms and signal processing normally needed to discern the pulse period and signal. In this paper, we describe an unrestricted pulse pressure sensor employing hetero-core fiber optics made of silicone rubber. We used the wrist's radial artery as the measurement site in subjects because biomedical monitoring at the wrist is suitable for natural daily activity in practical application settings, such as having the sensor on the armrest of a chair or strapped to the wrist as a wearable device. To detect minute pulse pressure changes from the radial artery at the wrist, we devised a fiber sensor arrangement such that minute pressure changes cause moderate bending of the three-point bending support of the sensing portion in the hetero-core fiber sensor. The proposed construction enabled the pulse sensor to achieve a high signal-to-noise ratio without complex signal processing to pick up the pulse signal. For unrestricted daily health management it is favorable that pulse pressure measurement can be achieved by simply placing a hand on the instrument without noticing the sensor. For unconstrained pulse pressure monitoring, two prototypes were developed and demonstrated: one was embedded in the base of an armrest and the other was a wearable device. 2. Fiber optic pulse pressure sensor 2.1 Hetero-core fiber optic sensor The proposed hetero-core optical fiber sensor consists of a 9 μm core SM transmission fiber and a 5 or 3 μm fiber (with lengths of a few mm), inserted using fusion splicing, as shown in Fig. 1. A hetero-core fiber with a 3 μm core fiber has been found to have relatively high optical loss sensitivity to the bending action of the hetero-core portion; however, it suffers large insertion losses of more than 10 dB. Conversely, the insertion loss of a 5 μm core hetero-core fiber has been found to be smaller; however, it has lower sensitivity to bending than the 3 μm core fiber. We employed a 1.31 μm wavelength fiber because our previous work has shown that the hetero-core fiber optic bending sensors have monotonic sensitivity at a wavelength of 1.31 μm [19]. Figure 2 panels (a) and (b) indicate the loss characteristics of the hetero-core fiber sensor with an approximated hetero-core portion curvature, with inserted hetero-core diameters of 5 and 3 μm, respectively. For the bending tests shown in Fig. 2, the hetero-core portion, located at the center of the two clamping points, which are separated by 35 mm, was bent at one clamping point, thus allowing for displacement toward the other clamping point.
Fig. 1. Single-mode-based hetero-core fiber optic bending sensor.
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Fig. 2. Bending loss characteristics of hetero-core fiber sensors; (a) inserted core diameter of 5 μm with hetero-core insertion length of 1.0, 1.5, and 2.0 mm and (b) inserted core diameter of 3 μm with hetero-core insertion length of 1.0 mm.
As shown in Fig. 2, it is apparent that the optical loss increases with increasing curvature. In addition, for a hetero-core diameter of 5 μm, Fig. 2(a), there is a tradeoff between bending sensitivity and the linearity of the insertion length of the hetero-core portion. In particular, the sensor with a 2 mm hetero-core insertion length can be used as a high-sensitivity sensor within a selective operation curvature. Therefore, for a monotonic property ofbending loss characteristics with curvature, the hetero-core fiber sensor needs to have a hetero-core insertion length less than 2.0 mm. The resulting insertion loss of less than 1 dB is small enough to warrant the use of a large signal-to-noise ratio. By contrast, as shown in Fig. 2(b), the bending loss of the hetero-core sensor with an inserted core of 3 μm in diameter is 11 dB for a curvature of 0.13 with monotonic property. The insertion loss owing to the 3 μm diameter hetero-core is relatively large at approximately 10 dB. In the case of the hetero-core 3-μm insertion length more than 1 mm, the insertion and bending loss might be large, as a result, the sensor has the insufficient signal-to-noise ratio. Therefore, initial bending loss resulting from initial settings such as cuff pressure should preferably be avoided. For a pulse sensor embedded in a living space, placing a wrist on the sensor will give the cuff pressure at the radial artery. Consequently, a fiber sensor with low bending sensitivity was able to detect pulse pressure changes. However, wearable sensing devices require relatively high bending sensitivity because it is difficult to maintain the cuff pressures. Therefore, we applied the 5 μm diameter hetero-core fiber sensor, which induces low insertion loss, to the living space monitors, while we used the 3 μm diameter ones, which have high sensitivity, in the wearable sensing devices. 2.2 Structure of fiber optic pulse pressure sensor Figure 3 shows the pulse pressure sensor structures based on hetero-core fiber optics. As shown in Fig. 3, to efficiently induce transformed bending in the hetero-core portion from the pulse pressure, three supporting points made of silicone rubber chips were placed at the center. Working the minute pulse pressure to the center of the chip induces large curvature bending of the hetero-core portion owing to the three supporting points in the pulse module. It was found
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Fig. 3. Pulse pressure-sensing module with hetero-core fiber optic sensors based on three-point bending made of silicone rubber: panels (a) and (b) show the structure of a pulse sensor with a hetero-core diameter of 5 μm; (c) shows the extended measurable area by means of multiple fiber sensors; and (d) shows a wearable minimized sensing device with a hetero-core diameter of 3μm.
in our previous work [18] that the structure of three supporting points using a hetero-core fiber optic sensor has monotonic characteristics with regard to weight change. This work indicated the force sensitivity of the structure of three supporting points. The tensile strength of silicone rubber is about 7 MPa. Figure 3 panels (a)–(c) show the structure of the pulse pressure module to be embedded in the living space. The radius of curvature of the heterocore portion in the pulse sensor module was 15 mm to allow for easy wiring of the fiber line. In addition, this allowed the sensor to have a high sensitivity because the plane containing the curved hetero-core fiber was perpendicular to the pulse pressure direction. The fiber line was fixed at two supporting points with a 90° circumferential angle. To estimate the sensitivity of the hetero-core portion at a relative position to the central support point, the position of the hetero-core portion was defined as the distance between the center of the hetero-core length and the central support position, P, as shown in Fig. 3(a). Conversely, to extend the detectable area in the pulse module for unconscious monitoring, three hetero-core sensors were arranged in parallel, as shown in Fig. 3(c). Figure 3(d) shows the structure of the pulse sensor as a wearable device. To minimize the size of the sensing device, the distance between the two supporting points was restricted to 8 mm. Consequently, changes in the curvature of the hetero-core portion are induced by pulse pressure. Because the embedded fiber sensor with three supporting points has resonant frequencies as high as a few kHz, the frequencies of the pulse pressure waveform, which generally peak at a few Hz, lie in a sufficiently flat part of the frequency spectrum. The distances between two supporting points can be changed to 7.1 or 2.4 mm per degree rise in temperature for two fixed distances of 23.6 and 8 mm, as shown in Fig. 3 panels (a) and (d), respectively; this is because the linear expansion coefficient of silicone rubber is about 300 × 10−6 °C−1. In addition, the hetero-core fiber sensor is in principle independent of temperature fluctuations, i.e., the proposed sensor is insusceptible to changes in ambient temperature. 2.3 Experimental setup Figure 4 shows the hetero-core pulse pressure sensor system. An optical power meter was used to measure the 1.3 μm wavelength light-emitting diode (LED) and laser diode (LD) light sources used for monitoring the 5 and 3 μm diameter hetero-core fibers, respectively. Although it was difficult to couple the LED to the SM fiber, the system has a sufficient
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signal-to-noise ratio owing to the 5 μm hetero-core fiber’s insertion and bending loss being up to a few dB. The LED and hetero-core fiber sensor system balances cost and system performance.
Fig. 4. Experimental setup for fiber optic pulse pressure monitoring; (a) measurement equipment and pulse pressure-sensing module set into a base and (b) minimized pulse pressure device on the wrist.
To evaluate the pulse sensor with cuff pressure at the wrist, the pulse sensor was located in the base, as shown in Fig. 4(a). Further, Fig. 4(b) shows the wearable arrangement, which exerts low cuff pressure. The minimized pulse sensor (shown in Fig. 3(d)) of the wearable device was attached to the wrist using a skin-compatible seal as shown in Fig. 4(b). 3. Experimental results and discussion 3.1 Analysis of pulse pressure waveform It has been demonstrated that a simple Fourier decomposition will allow the significant harmonics of the blood pressure pulse waveform to be observed [20]. Therefore, we evaluated the frequency and harmonic characteristics of the pulse waveform detected by the proposed pulse sensor. To evaluate the frequency characteristics of the time-varying pulse waveform, we used wavelet transformation analysis. Wavelet transformation is an attractive tool for data analysis in the field of biological signal processing [4] because the frequency responses of biological signals, such as pulse waveforms, tend to change over time. A wavelet transformation of a time-varying signal f(t) consists of computing the coefficients that are the inner products of f(t) against a wavelet series [11].These wavelets are labeled by the scale and time-shift location parameters a and b. In a continuous wavelet transform, the wavelet corresponding to scale a and time location b is Ψ a ,b (t ) =
t −b Ψ . a a
1
(1)
where Ψ(t) is the wavelet prototype. The continuous wavelet transform (CWT) is given by CWT { f (t ); a, b} = f (t )Ψ *a ,b (t )dt .
(2)
where the asterisk denotes complex conjugation. The wavelet coefficients of the CWT indicate the decomposed factors of these wavelets in the raw signal, f(t). This general inverse relationship between scale and frequency holds true for signals. Figure 5 displays the pulse pressure raw signal as the optical loss change from the hetero-core fiber pulse sensor and its wavelet from the decomposed waveforms. When the device was put on the wrist, the baseline of optical loss was defined to be 0 dB. As shown in Fig. 5(a), there are two pressure peaks during a single pulse. The pulse pressure waveform was sufficiently observed focusing on the only optical loss change. We chose the Gaussian function by taking the third derivative of the mother wavelet. Decomposed wavelet coefficients ranging from 23–26 are shown sequentially
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in Figs. 5(b)–5(e). The 25 scale wavelet-decomposed coefficient matches the basic frequency wave throughout the pulse waveform. The 24 and 25 coefficients contain similar peaks to the pulse waveform from the raw signal. The pseudo-frequency, Fc, corresponding to the scale a in Hz, is computed using the following relationship:
Fa =
Fc ., a∗Δ
(3)
where Δ is the sampling period and Fc is the center frequency of a wavelet in Hz. For a series of tests, the sampling period was 10 ms. Pseudo-frequencies were 5, 2.5, 1.25, and 0.625 Hz for the scale factors a of 23–26. The wavelet’s decomposed profile was temporally stable because the subject kept quiet during the test. Therefore, the basic frequency can be picked up by fast Fourier transform (FFT). Hence, band-pass filters, including the basic frequency around 1.25 Hz in the case of this pulse waveform, enable the basic frequency mode to be picked up. Pulsatile blood flow is directly connected to the pulse pressure oscillations between its systolic and diastolic values. The corresponding pressure difference as well as the absolute
Fig. 5. Wavelet decomposition of pulse pressure signal from the hetero-core fiber optic sensor: (a) real-time raw signal and (b)–(e) decomposed wavelet coefficients of the scale 23–26.
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Fig. 6. Pulse pressure waveform analysis based on FFT and criterion of pulse waveform amplitudes: (a) raw data of real-time loss profile for pulse sensor and its FFT spectrum; (b) band-pass-filtered spectrum to pick up first harmonic component; and (c) loss profile from inverse FFT of high-pass-filtered spectrum.
systolic and diastolic values represent important physiological parameters. To evaluate the pulse pressure waveform from the sensor data, the optical loss change corresponding to the pressure difference of the pulse waveform is estimated. The pressure difference corresponds to the amplitude of the pulse waveform's fundamental mode. Figure 6 illustrates the method used to extract the systolic and diastolic terms from the pulse waveform using adaptive bandpass and high-pass filters. Figure 6(a) shows the frequency response of the pulse waveform detected by the proposed pulse sensor. Harmonic components are shown at approximately 1.2, 2.2, and 3.1 Hz for the first, second, and third orders, respectively. First, the fundamental frequency of the pulse waveform is estimated from its FFT spectrum as shown in Fig. 6(a). The zero offset was varied to approximately 0.15 dB and the noise floor was lower than 0.01 dB, as shown in Fig. 6(a). Since band-pass filtering at 0.85 Hz–1.7 Hz can reduce motion artifacts and pick up the fundamental mode of the pulse waveform from the FFT spectrum in Fig. 6(a), the systolic and diastolic terms are extracted from the peaks of the waveform's fundamental mode, as shown in Fig. 6(b). The amplitude corresponding to the pulse pressure difference was estimated based on the peak terms using high-pass-filtered pulse waveforms, in which only the DC component was removed, as shown in Fig. 6(c). The amplitude of the band-pass-filtered waveform was temporally stable, which means the hysteresis could be almost neglected for the pulse measurement. Therefore, it is confirmed that an appropriate filter, such as a band-pass or high-pass filter, can pick up the systolic and diastolic terms as well as the pulse pressure difference.
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Fig. 7. Experimental results using a hetero-core fiber optic pulse-sensing module. Real-time responses in the loss for the position of the hetero-core portion of P = (a) 1 mm, (b) 0.5 mm, and (c) 0 mm.
Fig. 8. Loss amplitudes for the position of the hetero-core portion.
Fig. 9. Real-time loss profiles for expansion of the pulse-detectable area of the pulse pressure sensors (a) -A, (b) -B, and (c) -C, as shown in Fig. 3(c).
3.2 Sensitivity of fiber optic pulse sensor
Figure 7 shows the optical loss change real-time responses for different positions of the hetero-core portion, P, shown in Fig. 3(a). The pulse waveform amplitudes were defined in the aforementioned method based on fundamental frequency mode extraction during one
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pulse period, and were 0.24, 0.05, and 0.03 dB for the hetero-core positions of 0, 0.5, and 1.0 mm, respectively, as shown in Fig. 7 panels (a)–(c). Consequently, the amplitudes, which are pulse sensor sensitivities, decreased with increasing distance of the hetero-core position from the supporting center point. This is because the hetero-core portion length was 2 mm and the central spacer was 1 mm in width; therefore, the interfaces between the transmission and inserted fibers were efficiently given since the bending used the settings of the hetero-core portion on the central spacer. Figure 8 shows the amplitudes of the pulse waveforms and their deviations. For the 1.0 mm hetero-core position, the sensor was sometimes insensitive to the pulse pressure, which is represented by the amplitudes near 0 dB. Conversely, for all cases with the 0.5 and 0 mm hetero-core position, the sensor was sensitive to pulse pressures during multiple trials. The pulse pressure waveform was clearly detected by the proposed sensor as shown in Fig. 6; however, it was easily affected by the hetero-core position to the wrist. Therefore, we have not demonstrated a comparison with any gold standard pulse pressuresensing technology. The proposed device is a cost-effective system with low constraints for the subject. In addition, only the optical loss change sufficiently showed pulse pressure waveforms without calibration. The existing pulse monitoring technique using a plastic fiber optics has been presented [10]. The plastic fiber is based on multi-mode transmission, therefore, easily affected by the external disturbance such as bending and pressure. The proposed fiber optic pulse pressure sensor is made by the stable single-mode fiber, therefore, has robustness to the external disturbance and movement. Comparing to the existing fiber optic pulse sensor, the proposed fiber pulse sensor could be improved in reliability. Figure 9 shows the raw pulse pressure signals from the device containing the three hetero-core optic fiber sensors, which are arranged as in Fig. 3(c). The optical loss changes of the three heterocore optic fiber sensors -A, -B, and -C were simultaneously probed using a power meter with three channels. As shown in Fig. 9, focusing on the initial 9 s of the measurement, hetero-core sensor -C could detect pulses as shown in Fig. 9(c). Conversely, the other hetero-core sensors (-A and -B shown in Figs. 9(a) and 9(b)) had difficulty detecting the pulse pressure. However, after releasing the devices for a few seconds and putting them on the wrist again after 11s, the hetero-core sensor shown in Fig. 9(b) could detect a pulse signal. In contrast, the hetero-core sensor -C shown in Fig. 9(c) could no longer detect a pulse signal. In other words, pulse signals could be detected by different hetero-core sensors during each trial; therefore, changes in the configuration enable the detectable area to be enlarged. 3.3 Wearable pulse sensor tests
Figure 10 shows the pulse pressure waveform signals from the downsized module of a wearable fiber optic pulse sensor for subjects A, B, and C using the attachment shown in Fig. 4(b). To operate the pulse sensor under low-cuff-pressure conditions, a 3 μm core diameter fiber segment in a hetero-core portion was employed, which enabled the hetero-core fiber to have a higher sensitivity to the bending action. As shown in Fig. 10, all of the pulse waveform signals could be detected as optical loss waveforms with the low cuff pressures. The insertion loss of the hetero-core sensor increased compared with the hetero-core portion diameter of 5 μm; however, the signal-to-noise ratio was enough to measure the pulse waveforms. The amplitudes of the pulses were approximately 0.30, 0.14, and 0.07 dB for subjects A, B, and C, respectively; moreover, the noise floor was lower than 0.01 dB. The pulse waveforms for subject C in Fig. 10(b) were out-of-phase because the fiber sensor was bent in the initial state and straightened by the pulse pressure with inverse operation. Six subjects could measure their pulse using the wearable pulse module during multiple trials. The sensitivity can be adjusted by the combination in the core diameters of the hetero-core fiber. In order to increase the sensitivity, a small core diameter fiber, such as 3 μm, was employed at the hetero-core portion. The proposed fiber pulse pressure sensor uses a silica-fiber, whose diameter is 250 μm diameter including the coating layer for protection. The sensor module was 10 mm square as shown in Fig. 3(d), however, could be downsized because the fiber itself has small diameter such as 250 μm with flexibility.
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4. Conclusions
In this paper, we designed the characteristics of a practical hetero-core fiber optic pulse pressure sensor. The sensitivity of the proposed pulse pressure sensor was adjustable using the position of the hetero-core portion and the combination of core diameters in two spliced fibers. In practical tests of the hetero-core fiber sensor, multiple subjects used the pulse sensor
Fig. 10. Real-time responses of downsized and high-sensitive fiber optic pulse pressure sensor for the wearable device: subjects -A (a) -B (b), and -C (c).
without even noticing it. The wearable device was tested using the highly sensitive heterocore fiber sensors for unconstrained human activities. A measurement instrument including LEDs, optical power meters, and signal processing hardware, which can transmit the sensing signal through a wireless network, enables the proposed fiber pulse sensor to be portable. In conclusion, these results show that hetero-core optic sensors are useful for pulse pressure monitoring; in addition, the proposed sensor has several sensitivity advantages. These sensors could be useful for practical usage as a wearable device for health management in sports activities, as well as human health monitoring in the bathroom and at home.
Unconstrained pulse pressure monitoring for health management using hetero-core fiber optic sensor MICHIKO NISHIYAMA,* MASAKO SONOBE, AND KAZUHIRO WATANABE Department of Science and Engineering for Sustainable Innovation, Faculty of Science and Engineering, Soka university, 1-236 Tangi-Machi, Hachioji, Tokyo, 192-8577, Japan * [email protected]
Abstract: In this paper, we present a pulse pressure waveform sensor that does not constrain a wearer’s daily activity; the sensor uses hetero-core fiber optics. Hetero-core fiber sensors have been found to be sensitive to moderate bending. To detect minute pulse pressure changes from the radial artery at the wrist, we devised a fiber sensor arrangement using three-point bending supports. We analyzed and evaluated the measurement validity using wavelet transformation, which is well-suited for biological signal processing. It was confirmed that the detected pulse waveform had a fundamental mode frequency of around 1.25 Hz over the timevarying waveform. A band-pass filter with a range of frequencies from 0.85 to 1.7 Hz was used to pick up the fundamental mode. In addition, a high-pass filter with 0.85 Hz frequency eliminated arm motion artifacts; consequently, we achieved high signal-to-noise ratio. For unrestricted daily health management, it is desirable that pulse pressure monitoring can be achieved by simply placing a device on the hand without the sensor being noticed. Two types of arrangements were developed and demonstrated in which the pulse sensors were either embedded in a base, such as an armrest, or in a wearable device. A wearable device without cuff pressure using a sensitivity-enhanced fiber sensor was successfully achieved with a sensitivity of 0.07–0.3 dB with a noise floor lower than 0.01 dB or multiple subjects. ©2016 Optical Society of America OCIS codes: (060.2370) Fiber optics sensors; (170.3890) Medical optics instrumentation.
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#266540 Journal © 2016
http://dx.doi.org/10.1364/BOE.7.003675 Received 18 May 2016; revised 14 Aug 2016; accepted 22 Aug 2016; published 26 Aug 2016
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14. D. Lau, Z. Chen, J. T. Teo, S. H. Ng, H. Rumpel, Y. Lian, H. Yang, and P. L. Kei, “Intensity-modulated microbend fiber optic sensor for respiratory monitoring and gating during MRI,” IEEE Trans. Biomed. Eng. 60(9), 2655–2662 (2013). 15. L. Dziuda, F. W. Skibniewski, M. Krej, and J. Lewandowski, “Monitoring Respiration and Cardiac Activity Using Fiber Bragg Grating-Based Sensor,” IEEE Trans. Biomed. Eng. 59(7), 1934–1942 (2012). 16. M. Nishiyama, H. Sasaki, and K. Watanabe, “A deformation sensitive pad-structure embedded with hetero-core optic fiber sensors,” Sens. Actuators A Phys. 136(1), 205–211 (2007). 17. M. Nishiyama and K. Watanabe, “Wearable sensing glove with embedded hetero-core fiber-optic nerves for unconstrained hand motion capture,” IEEE Trans. Instrum. Meas. 58(12), 3995–4000 (2009). 18. M. Nishyama, M. Miyamoto, and K. Watanabe, “Respiration and body movement analysis during sleep in bed using hetero-core fiber optic pressure sensors without constraint to human activity,” J. Biomed. Opt. 16(1), 017002 (2011). 19. H. Sasaki, Y. Kubota, and K. Watanabe, “Sensitivity property of a hetero-core splice fiber optic displacement sensor,” Proc. SPIE 5579, 136–143 (2004). 20. J. K.-J. Li, Dynamics of the Vascular System (World scientific publishing, Singapore, 2004).
1. Introduction In recent years, the medical welfare network has advanced to an extent where diseases can be prevented before their occurrence [1–3]. Therefore, health monitoring techniques that people can perform themselves on a continuous basis are preferred, particularly, pulse pressure monitoring implemented for everyday activities such as sleeping and exercising [4–6]. These health monitoring devices should allow for continuous and autonomous detection of vital physiological indicators while being portable and comfortable from a user’s viewpoint. In other words, the user should not notice that they are wearing a sensor. In addition, the system must be robust to environmental disturbances such as electromagnetic interference (EMI), external temperature, and pressure fluctuations. Currently, there are several conventional methods for measuring someone’s pulse and blood pressure. However, an electrocardiogram (ECG) [7] or a photoelectric plethysmogram [8] may bring discomfort and inconvenience to the subject and physician because the sensors must be placed on the body. In addition, these techniques also require appropriate sensors to be installed on the subject. Conventional noninvasive blood flow pressure measurement is based on an inflatable pressure cuff surrounding an arm or a finger. However, periodic interruption of blood flow can affect a user’s physiological state and prevent comfortable sleep because this method requires inflation and deflation of the cuff. To meet the requirements for pulse measurement to be noninvasive and unconstrained, a low-cost pillow-shaped respiratory monitor [5] and a phonocardiographic sensor set on an air mat or a water mat that subjects sleep on [9] have been proposed. However, the raw signals generated using these unconstrained methods were noisy because of the inherent vibrational signals in the sensor output; consequently, an algorithm based on a wavelet transformation or appropriate band-pass filter was needed to be adapted to the raw signal. In the past few decades, several techniques have been developed for monitoring human respiration using optical fiber sensors, such as motion capture and monitoring systems [10–15]. This is because optical fiber sensors have several advantages, such as the sensor not needing a power supply, resistance to EMI, flexibility, light weight, and high resistance to moisture. In one example it was proposed to use plastic optical fibers for pulse oximetry [10]. However, plastic optical fibers have limited multi-mode stability for bending actions in the transmission line. Other approaches for respiration and heartbeat monitoring were employed, such as using optical interferometers with optical fiber length changes [13], microbend fiber sensors based on intensity modulation [14], and fiber Bragg gratings [15]. Although these techniques utilized an optical fiber, the optical interferometer methods were easily affected by external disturbances along the transmission fiber line because of the focus on the fiber length change owing to external perturbations. The microbend technique may induce overstrain in the fiber, and consequently, the initial transmission loss may seem to increase and the signal-to-noise ratio may decrease. Fiber Bragg gratings are well-known as strain and temperature sensors; however, they inherently need to compensate for the environmental temperature fluctuations for strain sensing. We have developed hetero-core optical fiber sensors [16–19] that can detect moderate bending with high sensitivity to the optical light intensity. The use of hetero-core optical fiber sensors
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has been proposed for displacement and pressure sensing in environmental monitoring, as well as for wearable motion analysis [17, 18]. This is because a hetero-core optical fiber sensor is sensitive only to the bending of the sensor portion. The fiber transmission line is unaffected by external disturbances such as pressure caused by the single-mode (SM) stable propagation scheme. Additionally, the fiber transmission line is independent of temperature fluctuations inside the sensor. Therefore, hetero-core optical fiber sensors have the potential to be used for unconstrained pulse pressure monitoring with high signal-to-noise ratio because of their stability to external disturbances. Such a health monitoring system can be simply constructed without the complex algorithms and signal processing normally needed to discern the pulse period and signal. In this paper, we describe an unrestricted pulse pressure sensor employing hetero-core fiber optics made of silicone rubber. We used the wrist's radial artery as the measurement site in subjects because biomedical monitoring at the wrist is suitable for natural daily activity in practical application settings, such as having the sensor on the armrest of a chair or strapped to the wrist as a wearable device. To detect minute pulse pressure changes from the radial artery at the wrist, we devised a fiber sensor arrangement such that minute pressure changes cause moderate bending of the three-point bending support of the sensing portion in the hetero-core fiber sensor. The proposed construction enabled the pulse sensor to achieve a high signal-to-noise ratio without complex signal processing to pick up the pulse signal. For unrestricted daily health management it is favorable that pulse pressure measurement can be achieved by simply placing a hand on the instrument without noticing the sensor. For unconstrained pulse pressure monitoring, two prototypes were developed and demonstrated: one was embedded in the base of an armrest and the other was a wearable device. 2. Fiber optic pulse pressure sensor 2.1 Hetero-core fiber optic sensor The proposed hetero-core optical fiber sensor consists of a 9 μm core SM transmission fiber and a 5 or 3 μm fiber (with lengths of a few mm), inserted using fusion splicing, as shown in Fig. 1. A hetero-core fiber with a 3 μm core fiber has been found to have relatively high optical loss sensitivity to the bending action of the hetero-core portion; however, it suffers large insertion losses of more than 10 dB. Conversely, the insertion loss of a 5 μm core hetero-core fiber has been found to be smaller; however, it has lower sensitivity to bending than the 3 μm core fiber. We employed a 1.31 μm wavelength fiber because our previous work has shown that the hetero-core fiber optic bending sensors have monotonic sensitivity at a wavelength of 1.31 μm [19]. Figure 2 panels (a) and (b) indicate the loss characteristics of the hetero-core fiber sensor with an approximated hetero-core portion curvature, with inserted hetero-core diameters of 5 and 3 μm, respectively. For the bending tests shown in Fig. 2, the hetero-core portion, located at the center of the two clamping points, which are separated by 35 mm, was bent at one clamping point, thus allowing for displacement toward the other clamping point.
Fig. 1. Single-mode-based hetero-core fiber optic bending sensor.
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Fig. 2. Bending loss characteristics of hetero-core fiber sensors; (a) inserted core diameter of 5 μm with hetero-core insertion length of 1.0, 1.5, and 2.0 mm and (b) inserted core diameter of 3 μm with hetero-core insertion length of 1.0 mm.
As shown in Fig. 2, it is apparent that the optical loss increases with increasing curvature. In addition, for a hetero-core diameter of 5 μm, Fig. 2(a), there is a tradeoff between bending sensitivity and the linearity of the insertion length of the hetero-core portion. In particular, the sensor with a 2 mm hetero-core insertion length can be used as a high-sensitivity sensor within a selective operation curvature. Therefore, for a monotonic property ofbending loss characteristics with curvature, the hetero-core fiber sensor needs to have a hetero-core insertion length less than 2.0 mm. The resulting insertion loss of less than 1 dB is small enough to warrant the use of a large signal-to-noise ratio. By contrast, as shown in Fig. 2(b), the bending loss of the hetero-core sensor with an inserted core of 3 μm in diameter is 11 dB for a curvature of 0.13 with monotonic property. The insertion loss owing to the 3 μm diameter hetero-core is relatively large at approximately 10 dB. In the case of the hetero-core 3-μm insertion length more than 1 mm, the insertion and bending loss might be large, as a result, the sensor has the insufficient signal-to-noise ratio. Therefore, initial bending loss resulting from initial settings such as cuff pressure should preferably be avoided. For a pulse sensor embedded in a living space, placing a wrist on the sensor will give the cuff pressure at the radial artery. Consequently, a fiber sensor with low bending sensitivity was able to detect pulse pressure changes. However, wearable sensing devices require relatively high bending sensitivity because it is difficult to maintain the cuff pressures. Therefore, we applied the 5 μm diameter hetero-core fiber sensor, which induces low insertion loss, to the living space monitors, while we used the 3 μm diameter ones, which have high sensitivity, in the wearable sensing devices. 2.2 Structure of fiber optic pulse pressure sensor Figure 3 shows the pulse pressure sensor structures based on hetero-core fiber optics. As shown in Fig. 3, to efficiently induce transformed bending in the hetero-core portion from the pulse pressure, three supporting points made of silicone rubber chips were placed at the center. Working the minute pulse pressure to the center of the chip induces large curvature bending of the hetero-core portion owing to the three supporting points in the pulse module. It was found
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Fig. 3. Pulse pressure-sensing module with hetero-core fiber optic sensors based on three-point bending made of silicone rubber: panels (a) and (b) show the structure of a pulse sensor with a hetero-core diameter of 5 μm; (c) shows the extended measurable area by means of multiple fiber sensors; and (d) shows a wearable minimized sensing device with a hetero-core diameter of 3μm.
in our previous work [18] that the structure of three supporting points using a hetero-core fiber optic sensor has monotonic characteristics with regard to weight change. This work indicated the force sensitivity of the structure of three supporting points. The tensile strength of silicone rubber is about 7 MPa. Figure 3 panels (a)–(c) show the structure of the pulse pressure module to be embedded in the living space. The radius of curvature of the heterocore portion in the pulse sensor module was 15 mm to allow for easy wiring of the fiber line. In addition, this allowed the sensor to have a high sensitivity because the plane containing the curved hetero-core fiber was perpendicular to the pulse pressure direction. The fiber line was fixed at two supporting points with a 90° circumferential angle. To estimate the sensitivity of the hetero-core portion at a relative position to the central support point, the position of the hetero-core portion was defined as the distance between the center of the hetero-core length and the central support position, P, as shown in Fig. 3(a). Conversely, to extend the detectable area in the pulse module for unconscious monitoring, three hetero-core sensors were arranged in parallel, as shown in Fig. 3(c). Figure 3(d) shows the structure of the pulse sensor as a wearable device. To minimize the size of the sensing device, the distance between the two supporting points was restricted to 8 mm. Consequently, changes in the curvature of the hetero-core portion are induced by pulse pressure. Because the embedded fiber sensor with three supporting points has resonant frequencies as high as a few kHz, the frequencies of the pulse pressure waveform, which generally peak at a few Hz, lie in a sufficiently flat part of the frequency spectrum. The distances between two supporting points can be changed to 7.1 or 2.4 mm per degree rise in temperature for two fixed distances of 23.6 and 8 mm, as shown in Fig. 3 panels (a) and (d), respectively; this is because the linear expansion coefficient of silicone rubber is about 300 × 10−6 °C−1. In addition, the hetero-core fiber sensor is in principle independent of temperature fluctuations, i.e., the proposed sensor is insusceptible to changes in ambient temperature. 2.3 Experimental setup Figure 4 shows the hetero-core pulse pressure sensor system. An optical power meter was used to measure the 1.3 μm wavelength light-emitting diode (LED) and laser diode (LD) light sources used for monitoring the 5 and 3 μm diameter hetero-core fibers, respectively. Although it was difficult to couple the LED to the SM fiber, the system has a sufficient
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signal-to-noise ratio owing to the 5 μm hetero-core fiber’s insertion and bending loss being up to a few dB. The LED and hetero-core fiber sensor system balances cost and system performance.
Fig. 4. Experimental setup for fiber optic pulse pressure monitoring; (a) measurement equipment and pulse pressure-sensing module set into a base and (b) minimized pulse pressure device on the wrist.
To evaluate the pulse sensor with cuff pressure at the wrist, the pulse sensor was located in the base, as shown in Fig. 4(a). Further, Fig. 4(b) shows the wearable arrangement, which exerts low cuff pressure. The minimized pulse sensor (shown in Fig. 3(d)) of the wearable device was attached to the wrist using a skin-compatible seal as shown in Fig. 4(b). 3. Experimental results and discussion 3.1 Analysis of pulse pressure waveform It has been demonstrated that a simple Fourier decomposition will allow the significant harmonics of the blood pressure pulse waveform to be observed [20]. Therefore, we evaluated the frequency and harmonic characteristics of the pulse waveform detected by the proposed pulse sensor. To evaluate the frequency characteristics of the time-varying pulse waveform, we used wavelet transformation analysis. Wavelet transformation is an attractive tool for data analysis in the field of biological signal processing [4] because the frequency responses of biological signals, such as pulse waveforms, tend to change over time. A wavelet transformation of a time-varying signal f(t) consists of computing the coefficients that are the inner products of f(t) against a wavelet series [11].These wavelets are labeled by the scale and time-shift location parameters a and b. In a continuous wavelet transform, the wavelet corresponding to scale a and time location b is Ψ a ,b (t ) =
t −b Ψ . a a
1
(1)
where Ψ(t) is the wavelet prototype. The continuous wavelet transform (CWT) is given by CWT { f (t ); a, b} = f (t )Ψ *a ,b (t )dt .
(2)
where the asterisk denotes complex conjugation. The wavelet coefficients of the CWT indicate the decomposed factors of these wavelets in the raw signal, f(t). This general inverse relationship between scale and frequency holds true for signals. Figure 5 displays the pulse pressure raw signal as the optical loss change from the hetero-core fiber pulse sensor and its wavelet from the decomposed waveforms. When the device was put on the wrist, the baseline of optical loss was defined to be 0 dB. As shown in Fig. 5(a), there are two pressure peaks during a single pulse. The pulse pressure waveform was sufficiently observed focusing on the only optical loss change. We chose the Gaussian function by taking the third derivative of the mother wavelet. Decomposed wavelet coefficients ranging from 23–26 are shown sequentially
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in Figs. 5(b)–5(e). The 25 scale wavelet-decomposed coefficient matches the basic frequency wave throughout the pulse waveform. The 24 and 25 coefficients contain similar peaks to the pulse waveform from the raw signal. The pseudo-frequency, Fc, corresponding to the scale a in Hz, is computed using the following relationship:
Fa =
Fc ., a∗Δ
(3)
where Δ is the sampling period and Fc is the center frequency of a wavelet in Hz. For a series of tests, the sampling period was 10 ms. Pseudo-frequencies were 5, 2.5, 1.25, and 0.625 Hz for the scale factors a of 23–26. The wavelet’s decomposed profile was temporally stable because the subject kept quiet during the test. Therefore, the basic frequency can be picked up by fast Fourier transform (FFT). Hence, band-pass filters, including the basic frequency around 1.25 Hz in the case of this pulse waveform, enable the basic frequency mode to be picked up. Pulsatile blood flow is directly connected to the pulse pressure oscillations between its systolic and diastolic values. The corresponding pressure difference as well as the absolute
Fig. 5. Wavelet decomposition of pulse pressure signal from the hetero-core fiber optic sensor: (a) real-time raw signal and (b)–(e) decomposed wavelet coefficients of the scale 23–26.
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Fig. 6. Pulse pressure waveform analysis based on FFT and criterion of pulse waveform amplitudes: (a) raw data of real-time loss profile for pulse sensor and its FFT spectrum; (b) band-pass-filtered spectrum to pick up first harmonic component; and (c) loss profile from inverse FFT of high-pass-filtered spectrum.
systolic and diastolic values represent important physiological parameters. To evaluate the pulse pressure waveform from the sensor data, the optical loss change corresponding to the pressure difference of the pulse waveform is estimated. The pressure difference corresponds to the amplitude of the pulse waveform's fundamental mode. Figure 6 illustrates the method used to extract the systolic and diastolic terms from the pulse waveform using adaptive bandpass and high-pass filters. Figure 6(a) shows the frequency response of the pulse waveform detected by the proposed pulse sensor. Harmonic components are shown at approximately 1.2, 2.2, and 3.1 Hz for the first, second, and third orders, respectively. First, the fundamental frequency of the pulse waveform is estimated from its FFT spectrum as shown in Fig. 6(a). The zero offset was varied to approximately 0.15 dB and the noise floor was lower than 0.01 dB, as shown in Fig. 6(a). Since band-pass filtering at 0.85 Hz–1.7 Hz can reduce motion artifacts and pick up the fundamental mode of the pulse waveform from the FFT spectrum in Fig. 6(a), the systolic and diastolic terms are extracted from the peaks of the waveform's fundamental mode, as shown in Fig. 6(b). The amplitude corresponding to the pulse pressure difference was estimated based on the peak terms using high-pass-filtered pulse waveforms, in which only the DC component was removed, as shown in Fig. 6(c). The amplitude of the band-pass-filtered waveform was temporally stable, which means the hysteresis could be almost neglected for the pulse measurement. Therefore, it is confirmed that an appropriate filter, such as a band-pass or high-pass filter, can pick up the systolic and diastolic terms as well as the pulse pressure difference.
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Fig. 7. Experimental results using a hetero-core fiber optic pulse-sensing module. Real-time responses in the loss for the position of the hetero-core portion of P = (a) 1 mm, (b) 0.5 mm, and (c) 0 mm.
Fig. 8. Loss amplitudes for the position of the hetero-core portion.
Fig. 9. Real-time loss profiles for expansion of the pulse-detectable area of the pulse pressure sensors (a) -A, (b) -B, and (c) -C, as shown in Fig. 3(c).
3.2 Sensitivity of fiber optic pulse sensor
Figure 7 shows the optical loss change real-time responses for different positions of the hetero-core portion, P, shown in Fig. 3(a). The pulse waveform amplitudes were defined in the aforementioned method based on fundamental frequency mode extraction during one
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pulse period, and were 0.24, 0.05, and 0.03 dB for the hetero-core positions of 0, 0.5, and 1.0 mm, respectively, as shown in Fig. 7 panels (a)–(c). Consequently, the amplitudes, which are pulse sensor sensitivities, decreased with increasing distance of the hetero-core position from the supporting center point. This is because the hetero-core portion length was 2 mm and the central spacer was 1 mm in width; therefore, the interfaces between the transmission and inserted fibers were efficiently given since the bending used the settings of the hetero-core portion on the central spacer. Figure 8 shows the amplitudes of the pulse waveforms and their deviations. For the 1.0 mm hetero-core position, the sensor was sometimes insensitive to the pulse pressure, which is represented by the amplitudes near 0 dB. Conversely, for all cases with the 0.5 and 0 mm hetero-core position, the sensor was sensitive to pulse pressures during multiple trials. The pulse pressure waveform was clearly detected by the proposed sensor as shown in Fig. 6; however, it was easily affected by the hetero-core position to the wrist. Therefore, we have not demonstrated a comparison with any gold standard pulse pressuresensing technology. The proposed device is a cost-effective system with low constraints for the subject. In addition, only the optical loss change sufficiently showed pulse pressure waveforms without calibration. The existing pulse monitoring technique using a plastic fiber optics has been presented [10]. The plastic fiber is based on multi-mode transmission, therefore, easily affected by the external disturbance such as bending and pressure. The proposed fiber optic pulse pressure sensor is made by the stable single-mode fiber, therefore, has robustness to the external disturbance and movement. Comparing to the existing fiber optic pulse sensor, the proposed fiber pulse sensor could be improved in reliability. Figure 9 shows the raw pulse pressure signals from the device containing the three hetero-core optic fiber sensors, which are arranged as in Fig. 3(c). The optical loss changes of the three heterocore optic fiber sensors -A, -B, and -C were simultaneously probed using a power meter with three channels. As shown in Fig. 9, focusing on the initial 9 s of the measurement, hetero-core sensor -C could detect pulses as shown in Fig. 9(c). Conversely, the other hetero-core sensors (-A and -B shown in Figs. 9(a) and 9(b)) had difficulty detecting the pulse pressure. However, after releasing the devices for a few seconds and putting them on the wrist again after 11s, the hetero-core sensor shown in Fig. 9(b) could detect a pulse signal. In contrast, the hetero-core sensor -C shown in Fig. 9(c) could no longer detect a pulse signal. In other words, pulse signals could be detected by different hetero-core sensors during each trial; therefore, changes in the configuration enable the detectable area to be enlarged. 3.3 Wearable pulse sensor tests
Figure 10 shows the pulse pressure waveform signals from the downsized module of a wearable fiber optic pulse sensor for subjects A, B, and C using the attachment shown in Fig. 4(b). To operate the pulse sensor under low-cuff-pressure conditions, a 3 μm core diameter fiber segment in a hetero-core portion was employed, which enabled the hetero-core fiber to have a higher sensitivity to the bending action. As shown in Fig. 10, all of the pulse waveform signals could be detected as optical loss waveforms with the low cuff pressures. The insertion loss of the hetero-core sensor increased compared with the hetero-core portion diameter of 5 μm; however, the signal-to-noise ratio was enough to measure the pulse waveforms. The amplitudes of the pulses were approximately 0.30, 0.14, and 0.07 dB for subjects A, B, and C, respectively; moreover, the noise floor was lower than 0.01 dB. The pulse waveforms for subject C in Fig. 10(b) were out-of-phase because the fiber sensor was bent in the initial state and straightened by the pulse pressure with inverse operation. Six subjects could measure their pulse using the wearable pulse module during multiple trials. The sensitivity can be adjusted by the combination in the core diameters of the hetero-core fiber. In order to increase the sensitivity, a small core diameter fiber, such as 3 μm, was employed at the hetero-core portion. The proposed fiber pulse pressure sensor uses a silica-fiber, whose diameter is 250 μm diameter including the coating layer for protection. The sensor module was 10 mm square as shown in Fig. 3(d), however, could be downsized because the fiber itself has small diameter such as 250 μm with flexibility.
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4. Conclusions
In this paper, we designed the characteristics of a practical hetero-core fiber optic pulse pressure sensor. The sensitivity of the proposed pulse pressure sensor was adjustable using the position of the hetero-core portion and the combination of core diameters in two spliced fibers. In practical tests of the hetero-core fiber sensor, multiple subjects used the pulse sensor
Fig. 10. Real-time responses of downsized and high-sensitive fiber optic pulse pressure sensor for the wearable device: subjects -A (a) -B (b), and -C (c).
without even noticing it. The wearable device was tested using the highly sensitive heterocore fiber sensors for unconstrained human activities. A measurement instrument including LEDs, optical power meters, and signal processing hardware, which can transmit the sensing signal through a wireless network, enables the proposed fiber pulse sensor to be portable. In conclusion, these results show that hetero-core optic sensors are useful for pulse pressure monitoring; in addition, the proposed sensor has several sensitivity advantages. These sensors could be useful for practical usage as a wearable device for health management in sports activities, as well as human health monitoring in the bathroom and at home.
