High-frequency Doppler Ultrasound Transducer for the Peripheral Circulatory System

Transcription

1 Journal of the Korean Physical Society, Vol. 59, No. 6, December 2011, pp High-frequency Doppler Ultrasound Transducer for the Peripheral Circulatory System Young Min Bae, Jeongwon Yang, Uk Kang and Guanghoon Kim Korea Electrotechnology Research Institute, Ansan , Korea (Received 15 December 2010) A Doppler ultrasound transducer was designed and implemented to measure the blood flow velocity in tiny vessels near the skin of hands or feet. The geometric parameters of the transducer for defining the observation volume were derived and implemented with an acoustic window made of polystyrene. The observation volume designed in this study was located 6.5 mm from the transducer, which was comparable to the value predicted geometrically. The two-way insertion loss of the transducer was 11.3 db on ultrasound frequency of 20 MHz, and the 3-dB bandwidth was approximately 2 MHz. In addition, the Doppler shift in the frequency measured by using a Doppler device composed of the transducer and a Doppler signal processing unit was proportional to the flow velocity generated by a homemade flowing system. Finally, we concluded that the transducer could be applied to measure the blood flow velocity in hands or feet. PACS numbers: p Keywords: Ultrasound, Doppler, Transducer, High-frequency, Peripheral circulatory system DOI: /jkps I. INTRODUCTION The cardiovascular system is a blood distribution network to transfer masses to and from cells in the body, and plays an important role in maintaining homeostasis. The function of the cardiovascular system is mainly regulated by the autonomic nervous system (ANS) that controls the visceral functions, and the dysfunction of it may bring about several vascular diseases [1]. In order to investigate the dysfunction or state of the cardiovascular system, parameters such as blood pressure or blood volume flow should be measured [2]. In particular, measuring the blood volume flows in vessels in hands or feet is useful in checking for abnormalities of the cardiovascular system [3]. For example, a neuropathy due to diabetes has been reported to result in an abnormality of blood flow in feet [4]. There are several methods to non-invasively measure the blood volume flow or linear velocity, such as photoplethysmographs (PPGs), laser Doppler methods, or ultrasound Doppler methods. The PPG provides only the change in the volume in the finger tips, and the laser Doppler method is limited to perfusion near the skin surface. On the contrary, as the Doppler ultrasound technique directly measures the change in the frequency of ultrasound reflected from objects such as red blood cells, it provides several advantages: (1) the blood flow velocity can be measured in the physical units such as centimeters per second, and (2) it can be applied to non-invasively measure the blood flow velocity of vessels inside the body. Indeed, Doppler ultrasound devices have been developed to measure the cerebral blood flow and the blood flow in the brachial artery [5,6]. As mentioned above, the change in the blood flow velocity in hands or feet is one potential biomarker for testing for dysfunctions of the cardiovascular system [7,8]. For example, as the blood flow velocity in hands or feet is changed against a stimulus such as a cold stimulus or deep breathing, measuring it is clinically valuable. However, as the vessels in hands or feet are located near the skin surface and the blood flow velocity in them is lower than the cerebral blood flow or the blood flow in the brachial artery, conventional Doppler devices cannot be applied to measure it. In this paper, the geometric parameters of a Doppler ultrasound transducer for measuring the blood flow velocity in hands or feet were determined, and a simple method for realizing the geometry of the transducer was implemented. In particular, the transducer implemented in this study has a dual-type structure in which the two piezoelectric materials are arrayed. The performances of the transducer were characterized in view of the two-way insertion loss as the sensitivity parameter. In addition, the feasibility of a Doppler ultrasound device composed of a transducer and a homemade Doppler signal processing unit was evaluated. kimbym@keri.re.kr; Tel:

2 High-frequency Doppler Ultrasound Transducer for the Peripheral Circulatory System Young Min Bae et al Table 1. Properties of the piezoelectric element. Value Composite PbZr 0.48Ti 0.52O 3 Shape Half disk (radius 3.25 mm) d C/N Electro-mechanical coupling coefficient (K t) 0.62 Fig. 1. Scheme of the dual-type Doppler ultrasound transducer. P1 and P2 are piezoelectric elements for transmitting and receiving ultrasound, respectively. d and a are the span and the angle between P1 and P2, respectively, and w is the length of the piezoelectric element. II. DESIGNS OF THE TRANSDUCER Fig. 2. (Color online) Structure of the transducer (not scaled). (a) structure of the transducer and (b) photograph of the acoustic window made of polystyrene. There are two types of Doppler ultrasound devices. One is operated in the continuous wave (CW) mode, and the other is operated in the pulse wave (PW) mode. In this study, for the CW mode, the Doppler ultrasound transducer was designed with a dual-type structure, in which two piezoelectric elements were arrayed, as shown in Fig. 1. Two piezoelectric elements (p1 and p2), in which one is for transmitting ultrasound and the other is for receiving ultrasound, are laid with some span (d) and angle (α). When the ultrasound rays transmitted and received are assumed to have the same width as the length of the piezoelectric element (w) and their direction is assumed to be perpendicular to the surface of piezoelectric elements in the longitudinal mode, the observation volume for measuring the blood flow velocity is defined as the area in which both the path of ultrasound transmitted and received are overlapped. In order to measure the blood velocity in the peripheral circulatory system such as hands or feet, the transducer should meet two requirements. First, the ultrasound frequency should be high enough to measure the low velocity, and second, the observation volume should be positioned near the skin. In this study, the ultrasound frequency was 20 MHz, higher than the frequencies (less than 10 MHz) of conventional Doppler transducers. In addition, the position of the observation volume was controlled by changing the geometry of the transducer. When the distance from the piezoelectric elements to the center of the observation volume, L, is defined in Fig. 1, the value of L can be calculated with Eq. (1), which shows that the value of L depends on the span (d) and angle (α): w + d cos α L = ( 2 α ). (1) 2 sin 2 By trial and error, the values of d and α were chosen as 0.5 mm and 20, respectively. Then the value of L was calculated to be 10.1 mm for a the piezoelectric element width of 3.25 mm. III. MATERIALS AND METHODS A PZT piezoelectric element, whose shape was half a disk with a radius of 3.25 mm, was purchased from Matsys Co. (Korea), and its main properties are shown in Table 1. An epoxy paste (EPO-TEK 301, Epoxy Technology, USA) to bond the piezoelectric elements to an acoustic window, silver paste (Dotite D-500, Fujikura Kasei, Japan), and silver wire (1 µm thick, California Fine Wire Company, USA) were purchased elsewhere. The structure of transducer is shown in Fig. 2(a). Using the epoxy paste, the piezoelectric elements were attached to the inner surface of an acoustic window, a photography of which is shown in Fig. 2(b). The wires signal1 and signal2 were bonded to the exposed surfaces of the piezoelectric elements by using the silver paste. The experimental setup was established to characterize the transducer, as shown in Fig. 3. It was composed of a function generator (33250A, Agilent Co., USA), a digital oscilloscope (WaveRunner 64Xi, LeCroy Co., USA), and silica glass (Thorlabs Inc., USA) as a reflector. The electric impedance of the cables and the connector used was 50 Ohm. The two-way insertion loss (IL) for estimating the sensitivity of the transducer was calculated with IL = 20 log P output P input + C, (2)

3 Journal of the Korean Physical Society, Vol. 59, No. 6, December 2011 Fig. 3. Experimental setup for characterizing the Doppler ultrasound transducer. Fig. 5. (Color online) Two-way insertion loss of transducer as a function of frequency as estimated by using the KLM model based a simulation. presented by the manufacturer, it was comparable to the conventional value of PZT materials [10]. 2. Performance of the transducer Fig. 4. (Color online) Electric impedance spectrum of the piezoelectric element. where P input is V p p (amplitude of the signal in voltage) generated by the function generator, P input is V p p measured by the oscilloscope, and C is compensation due to reflection from the silica glass (=0.793). IV. EXPERIMENTAL RESULTS 1. Characteristics of Piezoelectric Elements Figure 4 shows the electric impedance spectrum of the piezoelectric element measured by using the impedance analyzer (4194A, Agilent Co., USA). The thickness of the piezoelectric element was selected to be half the wavelength of the ultrasound in the piezoelectric element to get a center frequency of 20 MHz. The piezoelectric elements were 100µm thick, and the electro-mechanical coupling coefficient (K t ) was calculated with the resonance frequency and anti-resonance frequency, which are the minimum peak and the maximum peak in the electric impedance spectrum, respectively [9]. Although the calculated value of K t, 0.53, is lower than the nominal value The tone-burst response of the transducer was simulated with the KLM model, which is an equivalent electric circuit for a piezoelectric element [11]. In the model, the piezoelectric element which is represented by an electric transmission line and a perfect transformer, and an acoustic window, which is represented by the lossy electric transmission line, were connected in series. The thickness, acoustic impedance, and attenuation coefficient of polystyrene as the matching layer were 500 µm, 2.52 Mrayl, and 1.8 db/cm, respectively. The back and the front media of the piezoelectric element were set up to be air and water, respectively. The simulation was carried out using commercial software (PiezoCAD, Sonic Concept Co., USA). Figure 5 shows the IL of the transducer simulated as a function of the ultrasound frequency. Although the maximal value of IL was 7.5 db at 19.5 MHz, the 3-dB bandwidth was around 1 MHz, and the center frequency in the 3-dB bandwidth was 20 MHz, which is the designed value of the frequency. The value of IL was 8.5 db at a frequency of 20 MHz. The dual-type Doppler ultrasound transducer was fabricated according to the process described above. The piezoelectric element was attached to an acoustic window made of polystyrene. In the preliminary experiment, the attenuation coefficients of various polymers, such as PMMA (poly(methyl methacrylate)), polycarbonate, and polystyrene were measured, and the attenuation coefficient of polystyrene was 1.8 db/cm, lower than those of the others. Indeed, the acoustic window (Fig. 2(b)) of polystyrene was designed to keep the distance and the angle between the two piezoelectric elements as shown in Fig. 1.

4 High-frequency Doppler Ultrasound Transducer for the Peripheral Circulatory System Young Min Bae et al Fig. 7. (Color online) Doppler shift as a function of flow velocity. The inset is a typical Doppler sonogram of a branch of the palmar digital arteries. 3. Measurement of the Flow Velocity Fig. 6. (Color online) Characterization of the transducer in view of the two-way insertion loss. (a) change in the two-way insertion loss as a function of frequency in the real transducer, and (b) change in the two-way insertion loss as a function of distance between the transducer and the reflector. Figure 6(a) shows IL of the transducer as a function of the ultrasound frequency. The IL was calculated with Eq. (2) under the condition that a tone-burst signal with 20 cycles was applied to the transmitter of transducer. The frequency with the maximum IL and the bandwidth (3-dB width) were approximately 20.5 MHz and 2 MHz, respectively. The value of IL at 20 MHz was 11.3 MHz, which was similar to the value simulated with the KLM model. Figure 6(b) shows IL as a function of the distance between the transducer and the reflector. The distance from the transducer to the center of the observation volume can be controlled by changing the angle and the span between both the piezoelectric elements. In this study, the angle and the span were, respectively, selected as 0.5 mm and 20 to locate the center of the observation volume at approximately 10 mm from the transducer, for which the value of IL should be highest. From Fig. 6(b), the center of the observation volume was approximately located near 7.0 mm, lower than the designed value. The reason might be that as the shape of the piezoelectric element was round, the far end (L1 in Fig. 1) of the observation volume was closer to the transducer. The flow velocity was measured by using the transducer and a homemade Doppler signal processing unit with a heterodyne-based demodulator. The signal processing unit was composed of a signal generator of 20 MHz, an impedance matching circuit, a pre-amplifier with 60 db gain, and a heterodyne-based demodulator (MIQA-21D+, Minicircuit Co., USA) [12]. The demodulated signal, whose frequency is principally defined as a Doppler shift proportional to the flow velocity, was converted to a digital signal and was stored in a computer [12]. The system for generating the water flow was composed of a syringe pump and a silicon tube. Polystyrene beads (10 µm diameter) were diluted in water as reflectors mimicking red blood cells. The transducer and the silicon tube were contacted at an incident angle of 60 in a water bath filled with water. Figure 7 shows the Doppler shift as a function of the linear velocity of water flow. Although the change in the Doppler shift was negligible at flow velocities under m/s, it was linearly proportional to the flow velocity in the range of flow velocities from m/s to m/s, and the calculated slope was 22,700.0 Hz/m/s. As the blood flow velocity at arterioles in hands or feet is known to be m/s, the transducer can be applied to the Doppler ultrasound technique for non-invasively measuring the blood flow velocity in hands or feet [13]. The inset of Fig. 7 shows a typical Doppler sonogram of a branch of the palmar digital arteries in the top side of the finger, as measured using a Doppler device with a transducer. The horizontal and the vertical axes mean the time and the Doppler shift (in frequency), respectively. In the figure, the brightness represents the amplitude of the FFT spectrum of the Doppler shift (meaning the blood flow velocity) changed with heart beat. We confirmed that the transducer could be applied to measure the blood flow velocity in hands or feet.

5 Journal of the Korean Physical Society, Vol. 59, No. 6, December 2011 V. CONCLUSION It is clinically useful to measure the blood flow velocity in hands or feet. Here, we designed and built a dual-type Doppler transducer to non-invasively measure the blood flow velocity in vessels in hands or feet. Two geometric parameters, the span and the angle between both the piezoelectric elements were determined to reduce the distance between the transducer and the observation volume and were implemented with acoustic windows made of polystyrene. The position of the observation volume achieved with the acoustic window was comparable to the value calculated geometrically. In addition, the twoway insertion loss of transducer was 11.3 db at an ultrasound frequency of 20 MHz. In the experiment using a homemade flowing system, the Doppler shift in frequency measured by using the Doppler device with the transducer was observed to be linearly proportional to the flow velocities higher than m/s. In conclusion, the design of transducer and its implementation with the acoustic window can be used to develop a Doppler ultrasound device to measure the blood flow velocity in hands or feet. ACKNOWLEDGMENTS This work was supported by the Korea Foundation for International Cooperation of Science & Technology (KI- COS) through a grant provided by the Korean Ministry of Education, Science & Technology (MEST) in 2005 (No. K ). REFERENCES [1] J. V. Freeman, F. E. Dewey, D. M. Hadley, J. Myer and V. F. Froelicher, Prog. Cardiovasc. Dis. 48, 342 (2006). [2] P. Ganz and J. A. Vita, Circulation 108, 2049 (2003). [3] D. W. Bakker, IEEE Trans. Ultrason. Ferroelectr. Freq. Control 17, 170 (1970). [4] M. E. Edmonds, V. C. Roberts and P. J. Watkins, Diabetologia 22, 9 (1982). [5] M. A. Sloan et al., Neurology 62, 1468 (2004). [6] M. C. Corretti, G. D. Plotnick and R. A. Vogel, Am. J. Cardiol. 75, 783 (1995). [7] O. I. Kolev, Clin. Auton. Res. 13, 295 (2003). [8] S. V. Verma, M. R. Buchanan and T. J. Anderson, Circulation 108, 2054 (2003). [9] Q. Zhou et al., IEEE Trans. Ultrason. Ferroelectr. Freq. Control 54, 668 (2007). [10] F. S. Foster, L. K. Ryan and D. H. Turnbull, IEEE Trans. Ultrason. Ferroelectr. Freq. Control 38, 446 (1991). [11] R. Krimholtz, D. A. Leedom and G. L. Mattaei, Electron. Lett. 6, 398 (1970). [12] N. Aydin, L. Fan and D. H. Evans, Physiol. Meas. 15, 181 (1994). [13] V. A. Kozlov et al., Doppler Ultrasound Examination of Macro- and Microcirculatory Vessels of Neck, Face, and Mouth in Healthy and Some Pathological Conditions (SP minimax Ltd., St. Petersburg, 2000).