Annals of Biomedical Engineering (Ó 2014) DOI: 10.1007/s10439-014-1037-1
Microflotronic Arterial Tonometry for Continuous Wearable Non-Invasive Hemodynamic Monitoring PHILIP DIGIGLIO, RUYA LI, WENQI WANG, and TINGRUI PAN Micro-Nano Innovations (MiNI) Laboratory, Department of Biomedical Engineering, Department of Electrical and Computer Engineering, The University of California, Davis, CA 95616, USA (Received 3 March 2014; accepted 19 May 2014) Associate Editor Yong Xu oversaw the review of this article.
Abstract—Personalized mobile medicine will continue to advance through the development of wearable sensors that can wirelessly provide pertinent health information while remaining unobtrusive, comfortable, low cost, and easy to operate and interpret. It is the intention that the sensor presented hereafter can contribute to such innovation. By applying a combination of emerging microfluidic and electronic technologies, a miniature, flexible, transparent, highly sensitive and wearable pressure sensor with microfluidic elements has been implemented, referred to as a microflotronic device. High sensitivity of 0.1 kPa21 and fast response time on the order of tens of milliseconds has been achieved on the microflotronic sensor design. Its sensitivity is among the highest in impedance-based flexible pressure sensors. Once configured into an array, the transparent device can be easily aligned over the target artery to measure blood pressure noninvasively and continuously. In addition, the ultraflexible and thin plastic construct of the microflotronic sensor (of 270 lm in height) can be worn comfortably for extended periods of time. Importantly, the proposed microflotronic sensor has been utilized to perform arterial tonometry with the capability of noninvasive monitoring of arterial blood pressure waveforms in a real-time and continuous fashion. Keywords—Pressure sensors, Blood pressure, Microfluidics, Micro-electromechanical systems (MEMS), Cardiovascular disease (CVD), Arterial stiffness, Pulse, Pressure waveform, Mobile medicine, Wearable sensors.
Address correspondence to Tingrui Pan, Micro-Nano Innovations (MiNI) Laboratory, Department of Biomedical Engineering, Department of Electrical and Computer Engineering, The University of California, Davis, CA 95616, USA. Electronic mail: [email protected] Philip Digiglio and Ruya Li have contributed equally to this study.
INTRODUCTION Wearable health monitoring technologies have recently received enormous interest worldwide due to the rapidly aging global populations and the drastically increasing demand for in home healthcare.8 Body worn sensors, which can provide real-time continuous measurement of pertinent physiological parameters noninvasively and comfortably for extended periods of time, are of crucial importance for emerging applications of mobile medicine.8 Among those, analysis of blood pressure has been considered one of the most important clinical risk factors to detect, monitor, and predict cardiovascular disease (CVD),13 the leading cause of death worldwide.3,25 Early diagnosis of CVD can be critical in improving therapeutic outcomes. However, a major challenge with many forms of CVD is that the individual is often asymptomatic until the disease is highly progressed or until a possibly fatal event occurs.13,20 Therefore, wearable sensors and point-of-care devices that can be conveniently utilized in a home environment offer an appealing option for providing personalized health information to the doctors and health professionals.1 It is clinically accepted that invasive arterial and venous blood pressure monitoring systems can provide the most complete and accurate information concerning hemodynamic parameters.16,34 However, these systems can only be utilized in a surgical setting and impose the risks of infection and blood vessel injury.16 Therefore, noninvasive pressure measurement systems have been explored extensively over the past few decades. Doppler echocardiogram is a common approach utilized in a hospital setting.4,16 While this approach has the potential to provide fairly complete and accurate health information, results are highly dependent on the operation and interpretation performed by the trained medical Ó 2014 Biomedical Engineering Society
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professional.4,16 Another non-invasive approach that has recently gained popularity is arterial tonometry. It employs Pascal’s principle to measure the pressure in the artery based on the external force exerted on the artery wall over the compressed area, under the assumption that any confounding factors associated with the skin and tissue properties between the sensor and the artery wall have been largely ignored.35 This approach would simply allow a pressure sensor with the desired mechanical and electrical properties to continuously track the blood pressure waveform when aligned and pressed against the artery. Moreover, the pulsatile force measured by the tonometric sensor is frequently calibrated with a sphygmomanometer or an oscillometric device to determine the arterial pressure, as it is difficult to determine the contact area of the artery wall.19,34 Importantly, previous studies have shown that the arterial waveforms measured at various physiological locations can be utilized to quantitatively analyze pertinent hemodynamic parameters, which provide valuable diagnostic information about cardiovascular health and disease state, including, upstroke time (Tup),37 mean arterial pressure (MAP),29 central systolic blood pressure (cSBP),25 and carotidradial pulse wave velocity (PWV).18 Solid-state pressure transducers, based on either the piezoresistive or capacitive sensing principle, have been applied to arterial tonometry. Two prevalent device designs are a probe-shape and a watch-band.5,9,28 For instance, a flexible pulse monitoring system from PPS has been developed with twelve 2.5 mm 9 2.5 mm capacitive pressure sensing elements in a 3 by 4 matrix.9 However, its intrinsically limited capacitance makes its sensing performance vulnerable to environmental noise and parasitic capacitances. Another example is the Millar SPT device,5 which is composed of a piezoresistive pressure transducer implemented in a probe configuration. The device is compatible with common blood pressure monitoring systems, but external calibration must be performed with the measurement of the systolic and diastolic pressure of the brachial artery. As it is not fully integrated with its own monitoring system, data must often be post-processed in separate software, which makes real-time monitoring impractical.5 In another implementation, HealthStats’ CASPal is a watch-like arterial tonometry device calibrated with an oscillometric blood pressure cuff for home use. The rigid watch design monitors pressure changes in the calibrated systolic and diastolic blood pressure using a proprietary algorithm to estimate the cSBP. The main limitation of the device is its dependence on user placement, alignment and tension adjustment. Finally, the Tensys T-Line system offers a convenient wrist-mounted design with an extended data monitoring system. The device, utilizing a proprietary algorithm to calibrate the tonometric signals and to compute the radial systolic, diastolic, and MAP,
has been proven to have reasonably comparable accuracy to an invasive arterial catheter system when utilized in an Intensive Care Unit (ICU) setting.28 As the system is significantly prone to error from motion artifact, it is best utilized on patients that are under sedation or general anesthesia.28 In addition to its rigid construct, the device setup requires a medical professional to manually place the sensor on the patient’s wrist by following a cumbersome procedure. While current tonometric devices have successfully employed commercial pressure transducers to perform arterial pressure wave measurement, these systems generally cannot be worn comfortably for extended periods of time by an active individual, and moreover, it could be very challenging for any untrained individual to place and align the device over the optimal measurement site above the artery, from which considerable measurement errors can result. In this report, we have developed a miniature, flexible, transparent, wearable pressure sensor with highly sensitive microfluidic elements, referred to as a microflotronic device, by applying state-of-the-art microfluidics and electronics, for arterial blood pressure monitoring. The microflotronic sensor is intended to address the aforementioned challenges in the existing devices, by offering ultraflexibility for comfortable wear for extended periods of time, complete transparency for easier placement over the target area, and a redundant matrix of pressure sensing units to provide the operator with sufficient margin for alignment error. The microflotronic device represents the initial effort to incorporate a dynamic microfluidic layer as the sensing unit, compared to conventional solid-state designs. Specifically, sandwiched by two flexible transparent sensing membranes, the microfluidic layer experiences impedance changes as the membrane deforms in response to an externally applied pressure. By utilizing a low-viscosity sensing fluid, the microfluidic-based sensing structure allows the device to achieve not only the high device sensitivity of 0.1 kPa21, but also respond to mechanical stimuli with a response time on the order of tens of milliseconds. The high sensitivity of the sensor allows it to acquire a complete waveform signal even when the user only achieves partial applanation of the artery by applying minimal pressure over the artery. While remaining optically transparent and mechanically flexible, its response time is typically 10– 100 times faster than that of other flexible solid-state sensing counterparts,24,31 which is crucial for the accurate calculation of hemodynamic parameters, particularly, when the heart rate is elevated during exercise or mild tachycardia.36 Moreover, the flexible sensing system configured in a matrix format can be easily aligned over the radial artery and be worn comfortably for extended periods of time while performing arterial tonometry. This feature, in
Wearable Microflotronic Arterial Tonometry
FIGURE 1. Demonstration of the microflotronic pressure sensing array placed over the radial artery. When the microflotronic device is placed over the peripheral artery, tonometry can be performed by applying minimal pressure over the artery (note: the electrical connection to the device has been omitted from this concept photograph for clarity).
TABLE 1. Device feature comparison. Oscillometric cuff Heart rate Peripheral systolic blood pressure Peripheral diastolic blood pressure Mean arterial pressure Static measurement
Microflotronic sensor Heart rate Central systolic blood pressure Pulse wave velocity Mean arterial pressure Continuous measurement Upstroke time Complete waveform measurement
FIGURE 2. Illustration of the functional structure of the microflotronic device. It is composed of the microfluidic sensing layer of ethylene glycol encapsulated between the top and bottom flexible polyethylene terephthalate (PET) membranes, which are patterned with orthogonally aligned indium tin oxide (ITO) transparent electrodes arrays.
with orthogonally aligned indium tin oxide (ITO) transparent electrodes arrays. The microfluidic layer between the two sensing membranes is supported and separated by polymeric micropillars. Upon the microfluidic-electrode contact, electrical double layer (EDL) capacitance has formed.21,22 This configuration enables piezoresistive change of the microfluidic layer under the membrane deformation when a mechanical load is applied. To facilitate pressure mapping of the target arteries, a 3 by 3 microflotronic sensing array with 5 mm center-to-center spacing was designed and fabricated to achieve total surface coverage of 1.5 9 1.5 cm2. Fabrication Process
combination with the ultrahigh sensitivity of the microflotronic device, facilitates patient comfort during long-term continuous monitoring. The entire microflotronic device can be manufactured by industrycompatible microfabrication techniques, which allows massive production at a low unit cost in the foreseeable future.2,14 As a demonstration, the microflotronic pressure sensing array has been shown placed over the radial artery in Fig. 1. The features offered by the conventional oscillometric blood pressure cuff have been compared with the microflotronic sensor in Table 1.
MATERIALS AND METHODS Microflotronic Sensor Design Figure 2 illustrated the functional structure of the microflotronic pressure sensor. It is composed of the microfluidic sensing layer of ethylene glycol encapsulated between the top and bottom flexible polyethylene terephthalate (PET) membranes, which are patterned
The fabrication process started with a transparent PET film with 75 lm-thick ITO coating (Mianyang Prochema Commercial Co.), which was first lasermachined into the designated geometries as the top and bottom membranes of the device. Standard photolithography and wet etching processes were consecutively utilized to pattern ITO electrodes on the PET surface. The electrode arrays were designed to be parallel with the line width of 4.5 mm and the separation of 500 lm. After the completion of electrode patterning, the polymer micropillar layer of 120 lm in thickness was constructed between the electrode surfaces by laminating a negative dry-film photoresist (PerMX3050, DuPont) onto the bottom membrane.23 The two electrode membranes were aligned and subsequently packaged along the perimeter of the device utilizing a nanomolecular adhesion technique,7 while leaving the inlet and outlet ports for subsequent microfluidic loading. Following the assembly step, the sensing liquid, i.e., ethylene glycol (99.8%, SigmaArdrich) was loaded into the sandwiched microfluidic structure using the loading channels. The final device was again sealed by PDMS to prevent evaporation and
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environmental contamination. It is worth noting that the overlapped area between the two orthogonally aligned ITO electrodes formed the individual sensing region for each pressure sensing element, and the ITO wires also provided the electrical connection to the interfacial circuitry via a custom-made flexible printed circuit connector. Operation Principle The microfluidic layer encapsulated between the two ITO coated PET membranes is the sensing element of the microflotronic device. External pressure deforms the membrane causing resistive variation in the microfluidic layer. The relationship between the applied pressure and the resistance of the device can be modeled theoretically by analyzing the membrane deformation under external pressure and the corresponding microfluidic resistive variation.11 For pressure-induced membrane deformation, the membrane mechanical behavior can be well predicted by the classic thin-plate theory.6 This theory considers the contributions of radial and tangential stresses and strains to the membrane deformation.6 For the microfluidic resistive calculation, the classic Ohm’s law was employed utilizing a finite element approach.33 The final expression for the microfluidic layer resistance (R) can be expressed as: 0 R ¼ @r
Z
11 1 3 1 m2 a2 r2 5 þ v 2 2 T dSA a r P Eh3 16 1þv
S
ð1Þ where P is the pressure, a and h are the length and the thickness of the square sensing membrane respectively, E and m are the Young’s modulus and the Poisson ratio of the membrane material respectively, and r is the radial distance from the center of the sensing unit. T represents the original microfluidic layer thickness before deformation, whereas r is the electrical conductivity of the working fluid. S is the area of one sensing unit.
Device Characterization Device sensitivity was characterized by investigating the relationship between pressures applied onto device surface and the relative resistive change of the microflotronic sensing unit. The experimental characterization was conducted on the central unit of 3 by 3 microflotronic arrays. Three samples were prepared for the characterization. Each sensitivity measurement was repeated three times. The resistive measurement was directly assessed by a LCR meter (4284A, Agilent),
while a force gauge with 0.1 mN force resolution and 400 nm spatial resolution (M7-012, Mark-10) amounted to a motorized stage (VT-80, PImicos) provided the precise control of the loading force. By applying a concentrated load onto the center of the testing unit, pressure values were read out as the ratio of the force measurement from the gauge to the surface area of the sensing unit. Characterization of sensor response time was conducted by using a piezoelectric actuator (T215-A4-303 (Y), Piezo Systems, Inc.) which provided periodic movement (up to 50 Hz). The piezoelectric actuator was driven by a 10 V peak-to-peak square wave from 1 to 50 Hz for the device calibration. The corresponding resistive change of the sensing unit under the periodic mechanical stimuli of different frequencies was directly recorded at a sampling rate of 4080 Hz with a measurement system described hereafter. The entire measurement system consisted of the following components: a low voltage sinusoidal excitation signal provided by a waveform generator (Berkeley Nucleonics), the microflotronic sensing matrix described previously, custom-made flexible printed circuit connectors, an output amplifier in an inverting amplification mode (LM358AN op-amp, STMicroelectronics), an interfacial data acquisition system (USB-6210 Multifunction I/O, National Instruments) with LabVIEW software. The acquired measurement data were post-processed in MATLAB. To perform the measurement, a low excitation voltage of 1 V was applied to maximize the signal to noise ratio. Characterization of the optical transparency of the device was conducted by utilizing a Safire2 plate reader. A microflotronic array with a 100 lm thick microfluidic layer and 75 lm thick ITO/PET membranes was placed in the plate reader, and the transmission value of the device was read out after the plate reader scanned the visible light wavelengths. The measurement was conducted on 6 random spots on the device. Applications of Microflotronic Arterial Tonometry The entire microflotronic measurement system described above was utilized to perform arterial tonometry. The low excitation voltage prevented electrolysis while providing adequate signal-to-noise ratio. The operational resistance for each unit was 10 kX on average and 3 sensors in each row were excited sequentially leaving the overall power consumption less than 0.075 mW. An excitation and sampling frequency of 4080 Hz was utilized for the radial pulse measurement to provide sufficient frequency resolution to perform Fourier analysis. An excitation and sampling frequency of 1020 Hz was utilized for the carotid
Wearable Microflotronic Arterial Tonometry
pulse and carotid-radial pulse measurements. These frequencies that were multiples of 60 Hz were utilized in order to facilitate the use of a post-processing moving average filter to remove the power line noise. Fourier analysis was performed in order to confirm the successful measurement of the harmonic components of the pressure waveforms, which were crucial to reconstruct the original pressure waveforms. Applications of microflotronic arterial tonometry were demonstrated in this paper, including the measurement of the radial pressure waveform, the recording of the carotid pressure waveform, as well as the simultaneous acquisition of both the radial and the carotid pressure waveforms for assessing the carotid-radial PWV. For the arterial tonometric measurements, the identities of the healthy volunteers in the age group between 20 and 30 years old were undisclosed for the protection of their personal health information. All tonometric measurements were performed on the right side of the body with the volunteer seated. The right arm was comfortably supported during the radial tonometry experiments to ensure the radial artery was approximately in the same transverse plane as the heart. After the sensor was placed over the artery to be measured, one or more fingers were utilized to provide a gentle force to applanate the radial artery to perform tonometry measurements. The radial pulse measurement was externally calibrated to the diastolic and systolic blood pressure measured by an automatic blood pressure monitor (BP785, Omron). This device met the validation standards and feature recommendations for oscillometric devices provided by the American Heart Association.19,25 All hemodynamic parameters were averaged over a sample of 5 sequential waveforms.
likely due to the sensors being operated beyond the small deformation limits assumed in the theoretical model.6 The relative resistance of the microflontronic device with a 120 lm-thick sensing layer gradually decreased from 1 to 0.6 as the external pressure progressively elevated from 0 to 4 kPa. Whereas, for the 100 lm-thick sensing layer device, the relative resistance declined from 1 to 0.52 as the pressure was raised from 0 to 2.7 kPa. As it can be seen, the device sensitivity measurements coincided with the theoretical predictions presented in Eq. (1), which indicated that the importance of the microfluidic sensing layer in determining the device sensitivity and detection range. Curved Surface Performance Characterization The sensitivity characterization was repeated on curved surfaces with radii of 68 and 117 mm to characterize the performance of the flexible sensor within and beyond the normal range of the curvature of the underside of the wrist, which is 82–127 mm.10 Figure 3b plots the relative resistance R/R0 under various mechanical loads on the curved surfaces. The microflotronic devices characterized had a membrane thickness of 75 lm, a fluidic layer height of 100 lm and a spatial resolution of 5 mm. A gradual decrease in device sensitivity was observed with decreasing radii of curvature from 0.18 kPa21 on the flat surface, to 0.11 kPa21 on the 117 mm surface, to 0.09 kPa21. Importantly, the response remained linear on the curved surfaces such that performance variances due to wrist curvature would be resolved by the calibration with the automatic blood pressure monitor. Response Time Characterization
RESULTS AND DISCUSSION Sensitivity Characterization Device sensitivity of the microflotronic array was defined as the relative resistance change (DR/R0) vs. centralized pressure load (P). With a sensing membrane thickness of 75 lm and a microfluidic layer height of 120 lm, the microflotronic device 5 mm 9 5 mm unit area pressure sensor exhibited a sensitivity of 0.1 kPa21. As the microfluidic layer height reduced to 100 lm, the device sensitivity increased to 0.18 kPa21. Figure 3a summarized the characterization results (the dots with error bars) of the device sensitivity in both heights with the corresponding fitted curves (the dashed line) compared with the theoretical predictions (the solid line). The maximal divergence between theoretical predictions and experimental measurements is less than 5%. This is
Characterization of the response time of the microflotronic device with a 120 lm-thick sensing layer was shown in Figs. 3c–3e. As it can be seen, the relative resistance change of the device responded as the frequency of the mechanical stimuli was increased from 1 to 20 Hz, which indicated that the device can achieve a response time below 25 ms. The devices, in response to the mechanical stimuli, perform consistently at 1, 10 and 20 Hz pulsed mechanical loads at both high (0.96–1.00) and low (0.66–0.71) relative resistance range. It was important to note that the amplitude of the relative resistance measurement did not experience significant degradation as the frequency rose. This result meant the sensor behaved as a lowpass system with the pass band up to 20 Hz, which can be useful for rapid tracking of blood pressure signals. According to previous studies, a pass band of 20 Hz can be sufficient for accurately reproducing blood pressure waveforms up to 2 Hz or 120 beats per
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FIGURE 3. (a) Experimental and theoretical result of the 100 and 120 lm microfluidic layer devices sensitivities. The measurement results (dots) with error bars plotted with corresponding fitting curve (dashed lines) compare to sensitivity predicted by the theoretical model (solid lines). (b) Characterization of device sensitivity on surfaces of different radii of curvature. Mechanical responses of the device at two different load amplitudes under periodic stimuli of (c) 1 Hz, (d) 10 Hz and (e) 20 Hz.
minute.36 Such a frequency response requirement, based on Fourier analysis of blood pressure measurements, has indicated that the successful measurement of the first ten constituent harmonics of the primary pulse frequency would be sufficient to accurately reproduce the original blood pressure waveform.36 Arterial Waveform Acquisition The microflotronic array has been applied to conduct both radial and carotid arterial tonometry. In particular, radial arterial tonometry is a peripheral pulse measurement, whereas carotid arterial tonometry serves as a central pulse measurement. Figure 4 shows the successful reproduction of all key features of the pulse from both the radial and carotid arteries, which is crucial in order to accurately compute the associated
hemodynamic parameters.36 These key features include: (1) systolic upstroke, (2) systolic peak pressure, (3) reflected systolic peak pressure, (4) dicrotic notch, (5) diastolic runoff, and (6) end-diastolic pressure.15 Radial Arterial Pressure Mapping The microflotronic array configuration offered an automatic pulse tracking strategy for pulse recording which could not only record the pulse waveform in real-time, but also provided a dynamic pressure distribution over the skin surface that located the pulse position. Figure 5a showed the pressure distribution when the 3 by 3 microflotronic array device was worn against the radial artery on the wrist. Three recorded waveforms with different magnitudes provided examples of the signals acquired from 9 individual
Wearable Microflotronic Arterial Tonometry
FIGURE 4. (1) Systolic upstroke, (2) systolic peak pressure, (3) reflected systolic peak pressure, (4) dicrotic notch, (5) diastolic runoff, (6) end-diastolic pressure. The peripheral pulse was measured by placing the sensor over the radial artery of the wrist, and the central pulse was measured by placing the sensor over the carotid artery of the neck. The MAP calculated from the externally calibrated peripheral pulse was utilized to calibrate the central pulse along with the external diastolic pressure measurement.
pulse–recording sites. During the cardiovascular cycles, the consecutive pressure distribution maps were plotted at distinct physiological events in Figs. 5b–5d, respectively. As shown, the peak systolic pressure occurred at time t1 = 0.120 s, followed by the dicrotic notch at t2 = 0.341 s and the diastolic runoff at time t3 = 0.604 s. Furthermore, utilizing the microflotronic sensing array, the pressure distribution map can be of particular use to locate the position of the radial arterial pulse. The following advanced characteristics have been realized in the microflotronic device to provide key advantages for performing arterial tonometry. The microflotronic device has two key features to facilitate placement and alignment over the target pulse measurement site, which are identified as current challenges when utilizing commercial arterial tonometry systems. Firstly, the unprecedented transparency (greater than 80%) of the tonometric device facilitates easy placement over the target pulse measurement site. Secondly, the redundant matrix of pressure sensors featured in the pulse mapping demonstration provides the operator with significant margin for horizontal and vertical misalignment error while remaining largely insensitive to rotational misalignment error. The matrix design could readily be expanded to facilitate further misalignment tolerance, at which point the compromise between increased misalignment tolerance and increased demand for computing resources must be considered. The ultraflexibility of the microflotronic device allows sufficient comfortableness, particularly, when worn for extended periods of time for continuous monitoring. Furthermore, the PET in contact with the skin has been proven to cause no irritation in human skin patch tests.27 The main limitation of the flexibility of the device is the potential failure of the ITO electrodes under considerable tensile strains, which translates to a radius of curvature up to 17 mm.12 Together, the unique characteristics of transparency, matrix configuration and ultraflexibility permits the applica-
tion of the microflotronic device to a comfortable wearable arterial tonometry system that is convenient for home use on a mobile medicine platform, unlike current commercial arterial tonometry systems that are better suited for operation by trained medical professionals or use in a clinical setting. Hemodynamic Parameters Calculation The true value of continuous measurement of the arterial pressure waveform as opposed to static blood pressure measurements obtained from a sphygmomanometer or oscillometric cuff is that the entire waveform can be analyzed to yield a custom set of hemodynamic parameters that can provide CVD screening and personalized health information. One of the current challenges for hemodynamic analysis is the vast number of unique parameters that have been proposed and the lack of consensus on how to measure and compute these parameters.4,16,18,19,37 For this reason, a representative selection of hemodynamic parameters of proven clinical significance has been included in this paper. In particular, all peripheral hemodynamic parameters were calculated from the arterial tonometry experiments performed on the radial artery. As illustrated in Fig. 6, the peripheral systolic and diastolic blood pressure measurements from an automatic blood pressure monitor were utilized in a linear transformation to calibrate the blood pressure measurements from the radial tonometry experiments. The MAP values calculated from the radial tonometric data were utilized in a linear transformation to calibrate the blood pressure measurements from the carotid tonometry experiments. Specifically, the upstroke time (Tup) is a measurement of the duration of the systolic upstroke. It can be calculated as the time the previous end-diastolic pressure occurred subtracted from the time the systolic peak pressure occurs, as shown below.37 The reported
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FIGURE 5. Pressure mapping results of the radial artery. (a) Static pressure distribution on the 3 by 3 microflotronic array with three discrete radial artery waveforms recorded. During the cardiovascular cycles, the consecutive pressure distribution maps on the skin surface were plotted at distinct physiological events: (a) systolic pressure peak (t1), (b) dicrotic notch (t2), and (c) diastolic runoff (t3).
FIGURE 6. Illustration of (a) Peripheral calculation of the radial pulse pressure calibrated to the measured peripheral systolic and diastolic blood pressure (SBP and DBP). The difference of t2 and t1 is utilized to calculate the upstroke time (Tup). The arithmetic mean of the entire radial pulse is utilized to calculate the MAP. (b) Central calculation of cSBP as the peak pressure of the carotid pulse calibrated from the DBP and MAP of the radial pulse.
Tup distribution of 128.8 ± 9.8 ms compares reasonably with the normal distribution of 143 ± 52 ms of a control group of 17 individuals with a mean age of 74.6 years.37 Tup ¼ t2 t1 ;
ð2Þ
where t1 and t2 represent the beginning and the end of the systolic upstroke respectively. A slow upstroke time is an indicator of aortic stenosis. Clinical trials not only have proven a statistically significant correlation between increased upstroke time and arterial stenosis,
Wearable Microflotronic Arterial Tonometry
but also established a threshold upstroke time that can diagnose severe aortic stenosis with 93% sensitivity and 65% specificity.37 Moreover, the MAP is a key indicator of left ventricular contractility, vascular resistance, and vascular elasticity.29 MAP is independently a strong predictor of CVD, especially in men under age 60.29 The MAP was calculated as the arithmetic mean of the radial tonometric data. A MAP of 88 ± 10 mmHg was calculated from the radial tonometric data, and individual MAP values were utilized in the linear transform pressure calibration of the carotid waveforms. The reported MAP distribution reasonably agreed with the normal distribution of 93.0 ± 7.6 mmHg reported for a human study of 11,150 individuals with a mean age of 52.3 years.29 In fact, the increased risk for CVD is statistically significant for the 2nd quartile of the distribution of MAP for men below age 60.29 Thus, this is a pertinent physiological parameter to monitor to access CVD risk and general cardiovascular health. Importantly, the cSBP is an indicator of left ventricular function. cSBP has been proven to be a much stronger predictor of CVD than peripheral systolic blood pressure in individuals over 65 years of age.26 The cSBP was calculated as the average peak pressure of the calibrated carotid tonometric data. This calibration was possible as the diastolic pressure and MAP does not change significantly from the carotid artery to the radial artery.18,25 The reported cSBP distribution of 100 ± 12 had a close agreement with the normal range of cSBP of 101 ± 11 mmHg for the age range of 20 to 30 years reported by the International Journal of Angiology.30 cSBP measurements are also indifferent to confounding physiological factors that can have a significant impact on peripheral systolic blood pressure such as muscle contractions or local inflammation in the upper limb. Furthermore, carotid-radial PWV is a direct indicator of muscular artery stenosis (stiffness), because the transit path of the pulse traverses the arteries of the upper limb.18 It is strongly dependent on endothelial cell function. For this reason, a low-cost, simple, operator-independent test of dynamic endothelial cell function in response to ischemic conditions has been proposed utilizing carotid-radial PWV measurement.32 Importantly, the carotid-radial PWV can be calculated as the time delay (Dt) from the simultaneously measured end-diastolic pressure of the carotid waveform to the end-diastolic pressure of the radial waveform, that is, the carotid-radial PWV = L/Dt. It is important to note that post-process filtering with cutoff frequencies of 0.8 and 10 Hz was utilized to remove low and high frequency noise that could introduce inaccuracies in the determination of the temporal position of the enddiastolic pressure. As noted in Fig. 7, a 80 ms delay
FIGURE 7. The carotid-radial PWV is the speed at which the pulse travels from the carotid artery to the radial artery measured simultaneously at both sites. It is calculated as the path length over the time delay (Dt) of the end-diastolic pressure from the carotid artery to the radial artery.
TABLE 2. Measured arterial tonometric characteristics from healthy volunteers. Metric
Mean
SD
Units
MAP cSBP Tup
88.0 99.7 128.8
10.0 11.7 9.8
mmHg mmHg ms
was observed from the end-diastolic pressure of the carotid waveform to the end-diastolic pressure of the radial waveform. The path length of 69 cm was estimated using the body surface measurement from the suprasternal notch to the measurement site of the radial artery on the wrist of the volunteer.18 A carotidradial PWV of 8.6 m/s was computed from the simultaneous tonometric measurements. The calculated coratid-radial PWV of 8.6 m/s fell within the normal distribution of 9.01 ± 1.2 m/s reported for a healthy control group of 17 individuals with an age distribution of 32.8 ± 9.5 years.17 Table 2 summarizes the measured arterial tonometric characteristics from healthy young volunteers with an average age of 25 years old, with a mean peripheral systolic pressure of 111 mmHg and diastolic pressure of 74 mmHg.
CONCLUSIONS With the high transparency, ultraflexibility, and surface mapping capacity, the microflotronic sensing array has enabled a low-cost non-invasive hemodynamic monitoring solution that can be worn comfortably for extended periods of time without requiring a highly skilled operator to use. Such a system could
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advance personalized medicine by tracking a custom set of hemodynamic parameters in order to screen for CVD and provide personalized health information. A low-cost easy-to-use system suitable for point-of-care use could greatly increase access to hemodynamic analysis that could assist in the early diagnosis and prevention of heart disease. Furthermore, a robust and easily deployed system could contribute to the advancement of the study and standardization of hemodynamic parameters by making human studies more feasible.
ACKNOWLEDGMENTS This work is in part supported by the National Science Foundation (ECCS-0846502 and ECCS1307831) and the University of California Proof-ofConcept Program (PC-269200) to TP. RL acknowledges the fellowship support from China Scholarship Council (CSC). Authors would also like to thank Yijun Zhang for his assistance on the device illustrations.
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Bergmann, J. H. M., and A. H. McGregor. Body-worn sensor design: what do patients and clinicians want? Ann. Biomed. Eng. 39:2299–2312, 2011. 2 Bhowmik, A. K., Z. Li, and P. J. Bos. Mobile Displays: Technology and Applications. Chichester: Wiley, 2008. 3 World Health Organization. Cardiovascular diseases (CVDs). Fact sheet No. 317, 2012. http://www.who.int/ mediacentre/factsheets/fs317/en/. 4 Chew, M. S., and A. A˚neman. Hemodynamic monitoring using arterial waveform analysis. Curr. Opin. Crit. Care. 19:234–241, 2013. 5 Currie, K. D., M. Hubli, and A. V. Krassioukov. Applanation tonometry: a reliable technique to assess aortic pulse wave velocity in spinal cord injury. Spinal Cord 52(4):272–275, 2014. 6 Di Giovanni, M. Flat and Corrugated Diaphram Design Handbook. New York: Marcel Dekker Inc, 1982. 7 Ding, Y., S. Garland, M. Howland, A. Revzin, and T. Pan. Universal nanopatternable interfacial bonding. Adv Mater. 23(46):5551–5556, 2011. 8 Fraile, J. A., B. Javier, J. M. Corchado, and A. Abraham. Applying wearable solutions in dependent environments. IEEE Trans. Inf. Technol. Biomed. 14(6):1459–1467, 2010. 9 Hu, C. S., Y. F. Chung, C. C. Yeh, and C. H. Luo. Temporal and spatial properties of arterial pulsation measurement using pressure sensor array. Evid. Based Complement. Alternat. Med. 745127:1–9, 2012. 10 Keir, P. J., and R. P. Wells. Changes in geometry of the finger flexor tendons in the carpal tunnel with wrist posture and tendon load: an MRI study on normal wrists. Clin Biomech. 14:635–645, 1999.
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Li, R., B, Nie, P. Digiglio and T. Pan. Microflotronics: a flexible, transparent, pressure-sensitive microfluidic film. Adv. Funct. Mater., submitted, 2014. 12 Lin, H. K., S. M. Chiu, T. P. Cho, and J. C. Huang. Improved bending fatigue behavior of flexible PET/ITO film with thin metallic glass interlayer. Mater. Lett. 113:182– 185, 2013. 13 Lin, W. H., H. Zhang, and Y. T. Zhang. Investigation on cardiovascular risk prediction using physiological parameters. Comput. Math. Methods Med. 2013:272691, 2013. 14 Lueder, E. Liquid Crystal Displays (2nd ed.). United Kindom: A John Wiley and Sons, 2010. 15 Mark, J. B. Atlas of Cardiovascular Monitoring. New York: Churchill Livingstone, 1998. 16 Mateu Campos, M. L., A. Ferra´ndiz Selle´s, G. Gruartmoner de Vera, J. Mesquida Febrer, C. Sabatier Cloarec, Y. Poveda Herna´ndez, and X. Garcı´ a Nogales. Techniques available for hemodynamic monitoring. advantages and limitations. Med. Intensiva 36(6):434–444, 2012. 17 Mceleavy, O. D., R. W. Mccallum, J. R. Petrie, M. Small, J. M. C. Connell, N. Sattar, and S. J. Cleland. Higher carotidradial pulse wave velocity in healthy offspring of patients with type 2 diabetes.’’. Diabet. Med. 21(3):262–266, 2004. 18 Mitchell, G. F., S.-J. Hwang, R. S. Vasan, M. G. Larson, M. J. Pencina, N. M. Hamburg, J. A. Vita, D. Levy, and E. J. Benjamin. Arterial stiffness and cardiovascular events: the framingham heart study. Circulation. 121(4):505–511, 2010. 19 Munir, S., A. Guilcher, T. Kamalesh, B. Clapp, S. Redwood, M. Marber, and P. Chowienczyk. Peripheral augmentation index defines the relationship between central and peripheral pulse pressure. Hypertension. 51(1):112–118, 2007. 20 Naghavi, M., P. Libby, E. Falk, et al. From vulnerable plaque to vulnerable patient: a call for new definitions and risk assessment strategies: part II. Circulation 108(15):1772–1778, 2003. 21 Nie, B., R. Li, J. Brandt, and T. Pan. Iontronic microdroplet array for flexible ultrasensitive tactile sensing. Lab on a Chip. 14(6):1107–1116, 2014. 22 Nie, B., S. Xing, J. Brandt, and T. Pan. Droplet-based interfacial capacitive sensing. Lab on a Chip. 12:1110–1118, 2012. 23 Pan, T., and W. Wei. From cleanroom to desktop: emerging micro-nanofabrication technology for biomedical applications. Ann. Biomed. Eng. 39:600–620, 2010. 24 Pan, L., et al. An ultra-sensitive resistive pressure sensor based on hollow-sphere microstructure induced elasticity in conducting polymer film. Nat. Commun. 5:3002, 2014. 25 Pickering, T. G. Recommendations for blood pressure measurement in humans and experimental animals: part 1: blood pressure measurement in humans. Circulation 111(5):697–716, 2005. 26 Pini, R., M. C. Cavallini, V. Palmieri, et al. Central but not brachial blood pressure predicts cardiovascular events in an unselected geriatric population. J. Am. Coll. Cardiol. 51(25):2432–2439, 2008. 27 Polyethylene terephthalate. TOXNET Toxicology Data Network. Bethesda: National Institutes of Health, 2009. http://toxnet.nlm.nih.gov/cgi-bin/sis/search/a?dbs+hsdb:@ term+@DOCNO+7712. 28 Saugel, B., F. Fassio, A. Hapfelmeier, A. S. Meidert, R. M. Schmid, and W. Huber. The T-Line TL-200 system for continuous non-invasive blood pressure measurement in medical intensive care unit patients. Intensive Care Med. 38(9):1471–1477, 2012.
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Sesso, H. D., M. J. Stampfer, B. Rosner, C. H. Hennekens, J. M. Gaziano, J. E. Manson, and R. J. Glynn. Systolic and diastolic blood pressure, pulse pressure, and mean arterial pressure as predictors of cardiovascular disease risk in men. Hypertension. 36(5):801–807, 2000. 30 Sule, A. A., T. H. Hwang, and T. J. Chin. Very high central aortic systolic pressures in a young hypertensive patient on Telmisartan: is central aortic systolic pressure associated with white coat hypertension? Int. J. Angiol. 19(4):138–140, 2010. 31 Takei, K., et al. Nanowire active-matrix circuitry for lowvoltage macroscale artificial skin. Nat. Mater. 9(10):821– 826, 2010. 32 Torrado, J., D. Bia, Y. Zocalo, G. Valls, S. Lluberas, D. Craiem, et al. Reactive hyperemia-related changes in carotid-radial pulse wave velocity as a potential tool to characterize the endothelial dynamics. Conf. Proc. IEEE Eng. Med. Biol. Soc., pp. 1800–1803, 2009.
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Vassent, E., G. Meunier, A. Foggia, and G. Reyne. Simulation of induction of machine operation using a step by step finite-element method coupled with circuits and mechanical equations. IEEE Trans. Magn. 27(62):5232– 5234, 1991. 34 Webster, J. G., and J. W. Clark. Medical Instrumentation: Application and Design. Hoboken, NJ: Wiley, 2010. 35 Webster, J. G., and J. W. Clark. Blood pressure and sound. In: Medical Instrumentation: Application and Design, edited by J. G. Webster. Hoboken, NJ: Wiley, 2010, pp. 311–312. 36 Wilkinson, M. B., and M. Outram. Principles of pressure transducers, resonance, damping and frequency response. Anaesthesia Intensive Care Med. 10(2):102–105, 2009. 37 Yoshioka, N., Y. Fujita, T. Yasukawa, I. Sano, M. Kiso, M. Nakayama, et al. Do radial arterial pressure curves have diagnostic validity for identify severe aortic stenosis? J. Anesth. 24:7–10, 2010.
Microflotronic Arterial Tonometry for Continuous Wearable Non-Invasive Hemodynamic Monitoring PHILIP DIGIGLIO, RUYA LI, WENQI WANG, and TINGRUI PAN Micro-Nano Innovations (MiNI) Laboratory, Department of Biomedical Engineering, Department of Electrical and Computer Engineering, The University of California, Davis, CA 95616, USA (Received 3 March 2014; accepted 19 May 2014) Associate Editor Yong Xu oversaw the review of this article.
Abstract—Personalized mobile medicine will continue to advance through the development of wearable sensors that can wirelessly provide pertinent health information while remaining unobtrusive, comfortable, low cost, and easy to operate and interpret. It is the intention that the sensor presented hereafter can contribute to such innovation. By applying a combination of emerging microfluidic and electronic technologies, a miniature, flexible, transparent, highly sensitive and wearable pressure sensor with microfluidic elements has been implemented, referred to as a microflotronic device. High sensitivity of 0.1 kPa21 and fast response time on the order of tens of milliseconds has been achieved on the microflotronic sensor design. Its sensitivity is among the highest in impedance-based flexible pressure sensors. Once configured into an array, the transparent device can be easily aligned over the target artery to measure blood pressure noninvasively and continuously. In addition, the ultraflexible and thin plastic construct of the microflotronic sensor (of 270 lm in height) can be worn comfortably for extended periods of time. Importantly, the proposed microflotronic sensor has been utilized to perform arterial tonometry with the capability of noninvasive monitoring of arterial blood pressure waveforms in a real-time and continuous fashion. Keywords—Pressure sensors, Blood pressure, Microfluidics, Micro-electromechanical systems (MEMS), Cardiovascular disease (CVD), Arterial stiffness, Pulse, Pressure waveform, Mobile medicine, Wearable sensors.
Address correspondence to Tingrui Pan, Micro-Nano Innovations (MiNI) Laboratory, Department of Biomedical Engineering, Department of Electrical and Computer Engineering, The University of California, Davis, CA 95616, USA. Electronic mail: [email protected] Philip Digiglio and Ruya Li have contributed equally to this study.
INTRODUCTION Wearable health monitoring technologies have recently received enormous interest worldwide due to the rapidly aging global populations and the drastically increasing demand for in home healthcare.8 Body worn sensors, which can provide real-time continuous measurement of pertinent physiological parameters noninvasively and comfortably for extended periods of time, are of crucial importance for emerging applications of mobile medicine.8 Among those, analysis of blood pressure has been considered one of the most important clinical risk factors to detect, monitor, and predict cardiovascular disease (CVD),13 the leading cause of death worldwide.3,25 Early diagnosis of CVD can be critical in improving therapeutic outcomes. However, a major challenge with many forms of CVD is that the individual is often asymptomatic until the disease is highly progressed or until a possibly fatal event occurs.13,20 Therefore, wearable sensors and point-of-care devices that can be conveniently utilized in a home environment offer an appealing option for providing personalized health information to the doctors and health professionals.1 It is clinically accepted that invasive arterial and venous blood pressure monitoring systems can provide the most complete and accurate information concerning hemodynamic parameters.16,34 However, these systems can only be utilized in a surgical setting and impose the risks of infection and blood vessel injury.16 Therefore, noninvasive pressure measurement systems have been explored extensively over the past few decades. Doppler echocardiogram is a common approach utilized in a hospital setting.4,16 While this approach has the potential to provide fairly complete and accurate health information, results are highly dependent on the operation and interpretation performed by the trained medical Ó 2014 Biomedical Engineering Society
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professional.4,16 Another non-invasive approach that has recently gained popularity is arterial tonometry. It employs Pascal’s principle to measure the pressure in the artery based on the external force exerted on the artery wall over the compressed area, under the assumption that any confounding factors associated with the skin and tissue properties between the sensor and the artery wall have been largely ignored.35 This approach would simply allow a pressure sensor with the desired mechanical and electrical properties to continuously track the blood pressure waveform when aligned and pressed against the artery. Moreover, the pulsatile force measured by the tonometric sensor is frequently calibrated with a sphygmomanometer or an oscillometric device to determine the arterial pressure, as it is difficult to determine the contact area of the artery wall.19,34 Importantly, previous studies have shown that the arterial waveforms measured at various physiological locations can be utilized to quantitatively analyze pertinent hemodynamic parameters, which provide valuable diagnostic information about cardiovascular health and disease state, including, upstroke time (Tup),37 mean arterial pressure (MAP),29 central systolic blood pressure (cSBP),25 and carotidradial pulse wave velocity (PWV).18 Solid-state pressure transducers, based on either the piezoresistive or capacitive sensing principle, have been applied to arterial tonometry. Two prevalent device designs are a probe-shape and a watch-band.5,9,28 For instance, a flexible pulse monitoring system from PPS has been developed with twelve 2.5 mm 9 2.5 mm capacitive pressure sensing elements in a 3 by 4 matrix.9 However, its intrinsically limited capacitance makes its sensing performance vulnerable to environmental noise and parasitic capacitances. Another example is the Millar SPT device,5 which is composed of a piezoresistive pressure transducer implemented in a probe configuration. The device is compatible with common blood pressure monitoring systems, but external calibration must be performed with the measurement of the systolic and diastolic pressure of the brachial artery. As it is not fully integrated with its own monitoring system, data must often be post-processed in separate software, which makes real-time monitoring impractical.5 In another implementation, HealthStats’ CASPal is a watch-like arterial tonometry device calibrated with an oscillometric blood pressure cuff for home use. The rigid watch design monitors pressure changes in the calibrated systolic and diastolic blood pressure using a proprietary algorithm to estimate the cSBP. The main limitation of the device is its dependence on user placement, alignment and tension adjustment. Finally, the Tensys T-Line system offers a convenient wrist-mounted design with an extended data monitoring system. The device, utilizing a proprietary algorithm to calibrate the tonometric signals and to compute the radial systolic, diastolic, and MAP,
has been proven to have reasonably comparable accuracy to an invasive arterial catheter system when utilized in an Intensive Care Unit (ICU) setting.28 As the system is significantly prone to error from motion artifact, it is best utilized on patients that are under sedation or general anesthesia.28 In addition to its rigid construct, the device setup requires a medical professional to manually place the sensor on the patient’s wrist by following a cumbersome procedure. While current tonometric devices have successfully employed commercial pressure transducers to perform arterial pressure wave measurement, these systems generally cannot be worn comfortably for extended periods of time by an active individual, and moreover, it could be very challenging for any untrained individual to place and align the device over the optimal measurement site above the artery, from which considerable measurement errors can result. In this report, we have developed a miniature, flexible, transparent, wearable pressure sensor with highly sensitive microfluidic elements, referred to as a microflotronic device, by applying state-of-the-art microfluidics and electronics, for arterial blood pressure monitoring. The microflotronic sensor is intended to address the aforementioned challenges in the existing devices, by offering ultraflexibility for comfortable wear for extended periods of time, complete transparency for easier placement over the target area, and a redundant matrix of pressure sensing units to provide the operator with sufficient margin for alignment error. The microflotronic device represents the initial effort to incorporate a dynamic microfluidic layer as the sensing unit, compared to conventional solid-state designs. Specifically, sandwiched by two flexible transparent sensing membranes, the microfluidic layer experiences impedance changes as the membrane deforms in response to an externally applied pressure. By utilizing a low-viscosity sensing fluid, the microfluidic-based sensing structure allows the device to achieve not only the high device sensitivity of 0.1 kPa21, but also respond to mechanical stimuli with a response time on the order of tens of milliseconds. The high sensitivity of the sensor allows it to acquire a complete waveform signal even when the user only achieves partial applanation of the artery by applying minimal pressure over the artery. While remaining optically transparent and mechanically flexible, its response time is typically 10– 100 times faster than that of other flexible solid-state sensing counterparts,24,31 which is crucial for the accurate calculation of hemodynamic parameters, particularly, when the heart rate is elevated during exercise or mild tachycardia.36 Moreover, the flexible sensing system configured in a matrix format can be easily aligned over the radial artery and be worn comfortably for extended periods of time while performing arterial tonometry. This feature, in
Wearable Microflotronic Arterial Tonometry
FIGURE 1. Demonstration of the microflotronic pressure sensing array placed over the radial artery. When the microflotronic device is placed over the peripheral artery, tonometry can be performed by applying minimal pressure over the artery (note: the electrical connection to the device has been omitted from this concept photograph for clarity).
TABLE 1. Device feature comparison. Oscillometric cuff Heart rate Peripheral systolic blood pressure Peripheral diastolic blood pressure Mean arterial pressure Static measurement
Microflotronic sensor Heart rate Central systolic blood pressure Pulse wave velocity Mean arterial pressure Continuous measurement Upstroke time Complete waveform measurement
FIGURE 2. Illustration of the functional structure of the microflotronic device. It is composed of the microfluidic sensing layer of ethylene glycol encapsulated between the top and bottom flexible polyethylene terephthalate (PET) membranes, which are patterned with orthogonally aligned indium tin oxide (ITO) transparent electrodes arrays.
with orthogonally aligned indium tin oxide (ITO) transparent electrodes arrays. The microfluidic layer between the two sensing membranes is supported and separated by polymeric micropillars. Upon the microfluidic-electrode contact, electrical double layer (EDL) capacitance has formed.21,22 This configuration enables piezoresistive change of the microfluidic layer under the membrane deformation when a mechanical load is applied. To facilitate pressure mapping of the target arteries, a 3 by 3 microflotronic sensing array with 5 mm center-to-center spacing was designed and fabricated to achieve total surface coverage of 1.5 9 1.5 cm2. Fabrication Process
combination with the ultrahigh sensitivity of the microflotronic device, facilitates patient comfort during long-term continuous monitoring. The entire microflotronic device can be manufactured by industrycompatible microfabrication techniques, which allows massive production at a low unit cost in the foreseeable future.2,14 As a demonstration, the microflotronic pressure sensing array has been shown placed over the radial artery in Fig. 1. The features offered by the conventional oscillometric blood pressure cuff have been compared with the microflotronic sensor in Table 1.
MATERIALS AND METHODS Microflotronic Sensor Design Figure 2 illustrated the functional structure of the microflotronic pressure sensor. It is composed of the microfluidic sensing layer of ethylene glycol encapsulated between the top and bottom flexible polyethylene terephthalate (PET) membranes, which are patterned
The fabrication process started with a transparent PET film with 75 lm-thick ITO coating (Mianyang Prochema Commercial Co.), which was first lasermachined into the designated geometries as the top and bottom membranes of the device. Standard photolithography and wet etching processes were consecutively utilized to pattern ITO electrodes on the PET surface. The electrode arrays were designed to be parallel with the line width of 4.5 mm and the separation of 500 lm. After the completion of electrode patterning, the polymer micropillar layer of 120 lm in thickness was constructed between the electrode surfaces by laminating a negative dry-film photoresist (PerMX3050, DuPont) onto the bottom membrane.23 The two electrode membranes were aligned and subsequently packaged along the perimeter of the device utilizing a nanomolecular adhesion technique,7 while leaving the inlet and outlet ports for subsequent microfluidic loading. Following the assembly step, the sensing liquid, i.e., ethylene glycol (99.8%, SigmaArdrich) was loaded into the sandwiched microfluidic structure using the loading channels. The final device was again sealed by PDMS to prevent evaporation and
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environmental contamination. It is worth noting that the overlapped area between the two orthogonally aligned ITO electrodes formed the individual sensing region for each pressure sensing element, and the ITO wires also provided the electrical connection to the interfacial circuitry via a custom-made flexible printed circuit connector. Operation Principle The microfluidic layer encapsulated between the two ITO coated PET membranes is the sensing element of the microflotronic device. External pressure deforms the membrane causing resistive variation in the microfluidic layer. The relationship between the applied pressure and the resistance of the device can be modeled theoretically by analyzing the membrane deformation under external pressure and the corresponding microfluidic resistive variation.11 For pressure-induced membrane deformation, the membrane mechanical behavior can be well predicted by the classic thin-plate theory.6 This theory considers the contributions of radial and tangential stresses and strains to the membrane deformation.6 For the microfluidic resistive calculation, the classic Ohm’s law was employed utilizing a finite element approach.33 The final expression for the microfluidic layer resistance (R) can be expressed as: 0 R ¼ @r
Z
11 1 3 1 m2 a2 r2 5 þ v 2 2 T dSA a r P Eh3 16 1þv
S
ð1Þ where P is the pressure, a and h are the length and the thickness of the square sensing membrane respectively, E and m are the Young’s modulus and the Poisson ratio of the membrane material respectively, and r is the radial distance from the center of the sensing unit. T represents the original microfluidic layer thickness before deformation, whereas r is the electrical conductivity of the working fluid. S is the area of one sensing unit.
Device Characterization Device sensitivity was characterized by investigating the relationship between pressures applied onto device surface and the relative resistive change of the microflotronic sensing unit. The experimental characterization was conducted on the central unit of 3 by 3 microflotronic arrays. Three samples were prepared for the characterization. Each sensitivity measurement was repeated three times. The resistive measurement was directly assessed by a LCR meter (4284A, Agilent),
while a force gauge with 0.1 mN force resolution and 400 nm spatial resolution (M7-012, Mark-10) amounted to a motorized stage (VT-80, PImicos) provided the precise control of the loading force. By applying a concentrated load onto the center of the testing unit, pressure values were read out as the ratio of the force measurement from the gauge to the surface area of the sensing unit. Characterization of sensor response time was conducted by using a piezoelectric actuator (T215-A4-303 (Y), Piezo Systems, Inc.) which provided periodic movement (up to 50 Hz). The piezoelectric actuator was driven by a 10 V peak-to-peak square wave from 1 to 50 Hz for the device calibration. The corresponding resistive change of the sensing unit under the periodic mechanical stimuli of different frequencies was directly recorded at a sampling rate of 4080 Hz with a measurement system described hereafter. The entire measurement system consisted of the following components: a low voltage sinusoidal excitation signal provided by a waveform generator (Berkeley Nucleonics), the microflotronic sensing matrix described previously, custom-made flexible printed circuit connectors, an output amplifier in an inverting amplification mode (LM358AN op-amp, STMicroelectronics), an interfacial data acquisition system (USB-6210 Multifunction I/O, National Instruments) with LabVIEW software. The acquired measurement data were post-processed in MATLAB. To perform the measurement, a low excitation voltage of 1 V was applied to maximize the signal to noise ratio. Characterization of the optical transparency of the device was conducted by utilizing a Safire2 plate reader. A microflotronic array with a 100 lm thick microfluidic layer and 75 lm thick ITO/PET membranes was placed in the plate reader, and the transmission value of the device was read out after the plate reader scanned the visible light wavelengths. The measurement was conducted on 6 random spots on the device. Applications of Microflotronic Arterial Tonometry The entire microflotronic measurement system described above was utilized to perform arterial tonometry. The low excitation voltage prevented electrolysis while providing adequate signal-to-noise ratio. The operational resistance for each unit was 10 kX on average and 3 sensors in each row were excited sequentially leaving the overall power consumption less than 0.075 mW. An excitation and sampling frequency of 4080 Hz was utilized for the radial pulse measurement to provide sufficient frequency resolution to perform Fourier analysis. An excitation and sampling frequency of 1020 Hz was utilized for the carotid
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pulse and carotid-radial pulse measurements. These frequencies that were multiples of 60 Hz were utilized in order to facilitate the use of a post-processing moving average filter to remove the power line noise. Fourier analysis was performed in order to confirm the successful measurement of the harmonic components of the pressure waveforms, which were crucial to reconstruct the original pressure waveforms. Applications of microflotronic arterial tonometry were demonstrated in this paper, including the measurement of the radial pressure waveform, the recording of the carotid pressure waveform, as well as the simultaneous acquisition of both the radial and the carotid pressure waveforms for assessing the carotid-radial PWV. For the arterial tonometric measurements, the identities of the healthy volunteers in the age group between 20 and 30 years old were undisclosed for the protection of their personal health information. All tonometric measurements were performed on the right side of the body with the volunteer seated. The right arm was comfortably supported during the radial tonometry experiments to ensure the radial artery was approximately in the same transverse plane as the heart. After the sensor was placed over the artery to be measured, one or more fingers were utilized to provide a gentle force to applanate the radial artery to perform tonometry measurements. The radial pulse measurement was externally calibrated to the diastolic and systolic blood pressure measured by an automatic blood pressure monitor (BP785, Omron). This device met the validation standards and feature recommendations for oscillometric devices provided by the American Heart Association.19,25 All hemodynamic parameters were averaged over a sample of 5 sequential waveforms.
likely due to the sensors being operated beyond the small deformation limits assumed in the theoretical model.6 The relative resistance of the microflontronic device with a 120 lm-thick sensing layer gradually decreased from 1 to 0.6 as the external pressure progressively elevated from 0 to 4 kPa. Whereas, for the 100 lm-thick sensing layer device, the relative resistance declined from 1 to 0.52 as the pressure was raised from 0 to 2.7 kPa. As it can be seen, the device sensitivity measurements coincided with the theoretical predictions presented in Eq. (1), which indicated that the importance of the microfluidic sensing layer in determining the device sensitivity and detection range. Curved Surface Performance Characterization The sensitivity characterization was repeated on curved surfaces with radii of 68 and 117 mm to characterize the performance of the flexible sensor within and beyond the normal range of the curvature of the underside of the wrist, which is 82–127 mm.10 Figure 3b plots the relative resistance R/R0 under various mechanical loads on the curved surfaces. The microflotronic devices characterized had a membrane thickness of 75 lm, a fluidic layer height of 100 lm and a spatial resolution of 5 mm. A gradual decrease in device sensitivity was observed with decreasing radii of curvature from 0.18 kPa21 on the flat surface, to 0.11 kPa21 on the 117 mm surface, to 0.09 kPa21. Importantly, the response remained linear on the curved surfaces such that performance variances due to wrist curvature would be resolved by the calibration with the automatic blood pressure monitor. Response Time Characterization
RESULTS AND DISCUSSION Sensitivity Characterization Device sensitivity of the microflotronic array was defined as the relative resistance change (DR/R0) vs. centralized pressure load (P). With a sensing membrane thickness of 75 lm and a microfluidic layer height of 120 lm, the microflotronic device 5 mm 9 5 mm unit area pressure sensor exhibited a sensitivity of 0.1 kPa21. As the microfluidic layer height reduced to 100 lm, the device sensitivity increased to 0.18 kPa21. Figure 3a summarized the characterization results (the dots with error bars) of the device sensitivity in both heights with the corresponding fitted curves (the dashed line) compared with the theoretical predictions (the solid line). The maximal divergence between theoretical predictions and experimental measurements is less than 5%. This is
Characterization of the response time of the microflotronic device with a 120 lm-thick sensing layer was shown in Figs. 3c–3e. As it can be seen, the relative resistance change of the device responded as the frequency of the mechanical stimuli was increased from 1 to 20 Hz, which indicated that the device can achieve a response time below 25 ms. The devices, in response to the mechanical stimuli, perform consistently at 1, 10 and 20 Hz pulsed mechanical loads at both high (0.96–1.00) and low (0.66–0.71) relative resistance range. It was important to note that the amplitude of the relative resistance measurement did not experience significant degradation as the frequency rose. This result meant the sensor behaved as a lowpass system with the pass band up to 20 Hz, which can be useful for rapid tracking of blood pressure signals. According to previous studies, a pass band of 20 Hz can be sufficient for accurately reproducing blood pressure waveforms up to 2 Hz or 120 beats per
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FIGURE 3. (a) Experimental and theoretical result of the 100 and 120 lm microfluidic layer devices sensitivities. The measurement results (dots) with error bars plotted with corresponding fitting curve (dashed lines) compare to sensitivity predicted by the theoretical model (solid lines). (b) Characterization of device sensitivity on surfaces of different radii of curvature. Mechanical responses of the device at two different load amplitudes under periodic stimuli of (c) 1 Hz, (d) 10 Hz and (e) 20 Hz.
minute.36 Such a frequency response requirement, based on Fourier analysis of blood pressure measurements, has indicated that the successful measurement of the first ten constituent harmonics of the primary pulse frequency would be sufficient to accurately reproduce the original blood pressure waveform.36 Arterial Waveform Acquisition The microflotronic array has been applied to conduct both radial and carotid arterial tonometry. In particular, radial arterial tonometry is a peripheral pulse measurement, whereas carotid arterial tonometry serves as a central pulse measurement. Figure 4 shows the successful reproduction of all key features of the pulse from both the radial and carotid arteries, which is crucial in order to accurately compute the associated
hemodynamic parameters.36 These key features include: (1) systolic upstroke, (2) systolic peak pressure, (3) reflected systolic peak pressure, (4) dicrotic notch, (5) diastolic runoff, and (6) end-diastolic pressure.15 Radial Arterial Pressure Mapping The microflotronic array configuration offered an automatic pulse tracking strategy for pulse recording which could not only record the pulse waveform in real-time, but also provided a dynamic pressure distribution over the skin surface that located the pulse position. Figure 5a showed the pressure distribution when the 3 by 3 microflotronic array device was worn against the radial artery on the wrist. Three recorded waveforms with different magnitudes provided examples of the signals acquired from 9 individual
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FIGURE 4. (1) Systolic upstroke, (2) systolic peak pressure, (3) reflected systolic peak pressure, (4) dicrotic notch, (5) diastolic runoff, (6) end-diastolic pressure. The peripheral pulse was measured by placing the sensor over the radial artery of the wrist, and the central pulse was measured by placing the sensor over the carotid artery of the neck. The MAP calculated from the externally calibrated peripheral pulse was utilized to calibrate the central pulse along with the external diastolic pressure measurement.
pulse–recording sites. During the cardiovascular cycles, the consecutive pressure distribution maps were plotted at distinct physiological events in Figs. 5b–5d, respectively. As shown, the peak systolic pressure occurred at time t1 = 0.120 s, followed by the dicrotic notch at t2 = 0.341 s and the diastolic runoff at time t3 = 0.604 s. Furthermore, utilizing the microflotronic sensing array, the pressure distribution map can be of particular use to locate the position of the radial arterial pulse. The following advanced characteristics have been realized in the microflotronic device to provide key advantages for performing arterial tonometry. The microflotronic device has two key features to facilitate placement and alignment over the target pulse measurement site, which are identified as current challenges when utilizing commercial arterial tonometry systems. Firstly, the unprecedented transparency (greater than 80%) of the tonometric device facilitates easy placement over the target pulse measurement site. Secondly, the redundant matrix of pressure sensors featured in the pulse mapping demonstration provides the operator with significant margin for horizontal and vertical misalignment error while remaining largely insensitive to rotational misalignment error. The matrix design could readily be expanded to facilitate further misalignment tolerance, at which point the compromise between increased misalignment tolerance and increased demand for computing resources must be considered. The ultraflexibility of the microflotronic device allows sufficient comfortableness, particularly, when worn for extended periods of time for continuous monitoring. Furthermore, the PET in contact with the skin has been proven to cause no irritation in human skin patch tests.27 The main limitation of the flexibility of the device is the potential failure of the ITO electrodes under considerable tensile strains, which translates to a radius of curvature up to 17 mm.12 Together, the unique characteristics of transparency, matrix configuration and ultraflexibility permits the applica-
tion of the microflotronic device to a comfortable wearable arterial tonometry system that is convenient for home use on a mobile medicine platform, unlike current commercial arterial tonometry systems that are better suited for operation by trained medical professionals or use in a clinical setting. Hemodynamic Parameters Calculation The true value of continuous measurement of the arterial pressure waveform as opposed to static blood pressure measurements obtained from a sphygmomanometer or oscillometric cuff is that the entire waveform can be analyzed to yield a custom set of hemodynamic parameters that can provide CVD screening and personalized health information. One of the current challenges for hemodynamic analysis is the vast number of unique parameters that have been proposed and the lack of consensus on how to measure and compute these parameters.4,16,18,19,37 For this reason, a representative selection of hemodynamic parameters of proven clinical significance has been included in this paper. In particular, all peripheral hemodynamic parameters were calculated from the arterial tonometry experiments performed on the radial artery. As illustrated in Fig. 6, the peripheral systolic and diastolic blood pressure measurements from an automatic blood pressure monitor were utilized in a linear transformation to calibrate the blood pressure measurements from the radial tonometry experiments. The MAP values calculated from the radial tonometric data were utilized in a linear transformation to calibrate the blood pressure measurements from the carotid tonometry experiments. Specifically, the upstroke time (Tup) is a measurement of the duration of the systolic upstroke. It can be calculated as the time the previous end-diastolic pressure occurred subtracted from the time the systolic peak pressure occurs, as shown below.37 The reported
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FIGURE 5. Pressure mapping results of the radial artery. (a) Static pressure distribution on the 3 by 3 microflotronic array with three discrete radial artery waveforms recorded. During the cardiovascular cycles, the consecutive pressure distribution maps on the skin surface were plotted at distinct physiological events: (a) systolic pressure peak (t1), (b) dicrotic notch (t2), and (c) diastolic runoff (t3).
FIGURE 6. Illustration of (a) Peripheral calculation of the radial pulse pressure calibrated to the measured peripheral systolic and diastolic blood pressure (SBP and DBP). The difference of t2 and t1 is utilized to calculate the upstroke time (Tup). The arithmetic mean of the entire radial pulse is utilized to calculate the MAP. (b) Central calculation of cSBP as the peak pressure of the carotid pulse calibrated from the DBP and MAP of the radial pulse.
Tup distribution of 128.8 ± 9.8 ms compares reasonably with the normal distribution of 143 ± 52 ms of a control group of 17 individuals with a mean age of 74.6 years.37 Tup ¼ t2 t1 ;
ð2Þ
where t1 and t2 represent the beginning and the end of the systolic upstroke respectively. A slow upstroke time is an indicator of aortic stenosis. Clinical trials not only have proven a statistically significant correlation between increased upstroke time and arterial stenosis,
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but also established a threshold upstroke time that can diagnose severe aortic stenosis with 93% sensitivity and 65% specificity.37 Moreover, the MAP is a key indicator of left ventricular contractility, vascular resistance, and vascular elasticity.29 MAP is independently a strong predictor of CVD, especially in men under age 60.29 The MAP was calculated as the arithmetic mean of the radial tonometric data. A MAP of 88 ± 10 mmHg was calculated from the radial tonometric data, and individual MAP values were utilized in the linear transform pressure calibration of the carotid waveforms. The reported MAP distribution reasonably agreed with the normal distribution of 93.0 ± 7.6 mmHg reported for a human study of 11,150 individuals with a mean age of 52.3 years.29 In fact, the increased risk for CVD is statistically significant for the 2nd quartile of the distribution of MAP for men below age 60.29 Thus, this is a pertinent physiological parameter to monitor to access CVD risk and general cardiovascular health. Importantly, the cSBP is an indicator of left ventricular function. cSBP has been proven to be a much stronger predictor of CVD than peripheral systolic blood pressure in individuals over 65 years of age.26 The cSBP was calculated as the average peak pressure of the calibrated carotid tonometric data. This calibration was possible as the diastolic pressure and MAP does not change significantly from the carotid artery to the radial artery.18,25 The reported cSBP distribution of 100 ± 12 had a close agreement with the normal range of cSBP of 101 ± 11 mmHg for the age range of 20 to 30 years reported by the International Journal of Angiology.30 cSBP measurements are also indifferent to confounding physiological factors that can have a significant impact on peripheral systolic blood pressure such as muscle contractions or local inflammation in the upper limb. Furthermore, carotid-radial PWV is a direct indicator of muscular artery stenosis (stiffness), because the transit path of the pulse traverses the arteries of the upper limb.18 It is strongly dependent on endothelial cell function. For this reason, a low-cost, simple, operator-independent test of dynamic endothelial cell function in response to ischemic conditions has been proposed utilizing carotid-radial PWV measurement.32 Importantly, the carotid-radial PWV can be calculated as the time delay (Dt) from the simultaneously measured end-diastolic pressure of the carotid waveform to the end-diastolic pressure of the radial waveform, that is, the carotid-radial PWV = L/Dt. It is important to note that post-process filtering with cutoff frequencies of 0.8 and 10 Hz was utilized to remove low and high frequency noise that could introduce inaccuracies in the determination of the temporal position of the enddiastolic pressure. As noted in Fig. 7, a 80 ms delay
FIGURE 7. The carotid-radial PWV is the speed at which the pulse travels from the carotid artery to the radial artery measured simultaneously at both sites. It is calculated as the path length over the time delay (Dt) of the end-diastolic pressure from the carotid artery to the radial artery.
TABLE 2. Measured arterial tonometric characteristics from healthy volunteers. Metric
Mean
SD
Units
MAP cSBP Tup
88.0 99.7 128.8
10.0 11.7 9.8
mmHg mmHg ms
was observed from the end-diastolic pressure of the carotid waveform to the end-diastolic pressure of the radial waveform. The path length of 69 cm was estimated using the body surface measurement from the suprasternal notch to the measurement site of the radial artery on the wrist of the volunteer.18 A carotidradial PWV of 8.6 m/s was computed from the simultaneous tonometric measurements. The calculated coratid-radial PWV of 8.6 m/s fell within the normal distribution of 9.01 ± 1.2 m/s reported for a healthy control group of 17 individuals with an age distribution of 32.8 ± 9.5 years.17 Table 2 summarizes the measured arterial tonometric characteristics from healthy young volunteers with an average age of 25 years old, with a mean peripheral systolic pressure of 111 mmHg and diastolic pressure of 74 mmHg.
CONCLUSIONS With the high transparency, ultraflexibility, and surface mapping capacity, the microflotronic sensing array has enabled a low-cost non-invasive hemodynamic monitoring solution that can be worn comfortably for extended periods of time without requiring a highly skilled operator to use. Such a system could
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advance personalized medicine by tracking a custom set of hemodynamic parameters in order to screen for CVD and provide personalized health information. A low-cost easy-to-use system suitable for point-of-care use could greatly increase access to hemodynamic analysis that could assist in the early diagnosis and prevention of heart disease. Furthermore, a robust and easily deployed system could contribute to the advancement of the study and standardization of hemodynamic parameters by making human studies more feasible.
ACKNOWLEDGMENTS This work is in part supported by the National Science Foundation (ECCS-0846502 and ECCS1307831) and the University of California Proof-ofConcept Program (PC-269200) to TP. RL acknowledges the fellowship support from China Scholarship Council (CSC). Authors would also like to thank Yijun Zhang for his assistance on the device illustrations.
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