Novel Phased Array Piezoelectric Micromachined Ultrasound Transducers (pMUTs) for Medical Imaging

Two kinds of 128 channels pMUT-based phased array ultrasound transducers are described in this paper, one with a high frequency of around 6 MHz and another with a low frequency of around 1.5 MHz. The active area of the transducer is around 25 mm long and 10 mm wide. There are in total 6270 pMUT elements in the reported arrays. For the transducer with a center frequency of 6 MHz, each pMUT has a membrane diameter of 85 <inline-formula> <tex-math notation="LaTeX">$\mu \text{m}$ </tex-math></inline-formula> and the pitch between every two channels is 200 <inline-formula> <tex-math notation="LaTeX">$\mu \text{m}$ </tex-math></inline-formula>. The transducer benefits from a high transmission and receive sensitivity of 44 kPa/V/Channel @ 3 cm and 204 mV/MPa, respectively. For the transducer with a center frequency of 1.5 MHz, each pMUT has a membrane diameter of <inline-formula> <tex-math notation="LaTeX">$160~\mu \text{m}$ </tex-math></inline-formula> and the pitch between every two elements is 214 <inline-formula> <tex-math notation="LaTeX">$\mu \text{m}$ </tex-math></inline-formula>. The proposed transducer obtained a transmit and receive sensitivity of 430 Pa/V/Channel @ 3 cm and 190 mV/MPa, respectively. The transducer has a −3dB and −6dB bandwidth of 118% and 184%, respectively. The bandwidth is higher than any previously reported transducer in any technology, such as bulk-PZT, pMUT, or capacitive MUT (cMUT). The functionality of the transducer arrays is confirmed by obtaining B-mode images in water medium.

dicing of the bulk layer by a saw, which limits the pitch by the 35 kerf of the saw [9]. Also, the transducers use high actuation 36 voltages in the range of 70-140 V during transmission, which 37 limits their usage or performance in a compact, battery-based 38 devices. Moreover, the manufacturing of conventional trans-39 ducers is very costly and labor-intensive. 40 To overcome the aforementioned drawbacks and align 41 ultrasound imaging systems with the market trend of 42 electronic devices, micromachined ultrasound transducers 43 (MUTs) have been developed in the last decade. MUTs 44 benefit from a lithography-based MEMS fabrication process, 45 ultra-miniaturization, potential integration with CMOS tech-46 nology, and a low-cost fabrication process. There are two 47 types of micromachined technologies, namely piezoelectric 48 micromachined ultrasound transducer (pMUT) with a piezo-49 electric thin film typically in d31 mode [10], [11] and capac-50 itive micromachined ultrasound transducer (cMUT) also in 51 the flexural mode with electrostatic forces [12], [13], [14]. 52 A cMUT element is, in essence, a miniaturized capaci-53 tor that consists of a thin metalized suspended membrane. 54 sion response with respect to bulk piezoelectric technology. 88 In this paper, we propose 1D pMUT phased arrays with 89 128 channels to address these issues. The design and fabri-90 cation process of the pMUT arrays are explained in detail. 91 Moreover, measurement results of imaging prototypes are 92 described, which illustrate the promising potential of the 93 transducer for medical imaging applications. This paper is an 94 extension to our conference paper [17], and beyond what was  power. It has been previously demonstrated that in order to 106 have sufficient output acoustic power, the lateral dimension 107 of the pMUT membrane should be equal to or bigger than 108 half of the transmitted wavelength [16]. Therefore, based on 109 the desired working frequency, lateral dimensions and the 110 thickness of the membrane are optimized. On the other hand, 111 the bandwidth of each pMUT can be tuned by utilizing the 112 damping effect of the medium on the pMUT membrane. The 113 mass of the medium damps the membrane vibration more 114 effectively when the lateral dimension of the membrane is 115 smaller than the wavelength. Therefore, the lateral dimension 116 of the membrane should be a compromise between the output 117 acoustic power and the bandwidth. Due to the inverse square 118 relationship of frequency to aperture, a higher resonant fre-119 quency is likely to have a higher aperture to wavelength ratio 120 and hence higher acoustic power and transmission response. 121 In contrast, low frequency designs normally have a lower 122 aperture-wavelength ratio and are exposed to higher damping 123 resulting in higher bandwidths. Based on these aspects, two 124 designs are proposed in this work.

125
For the following discussion, the pMUT array is considered 126 to be in the x-z plane with the y-axis perpendicular to the 127 surface of the array. The azimuth angle is placed in the 128 x-y plane and any azimuth and elevation angle on the y-axis 129 is zero. The elements are placed along the azimuth x-axis 130 and each channel is elongated along the elevation direction 131 (z-axis). Since there is no significant beamforming in the 132 elevation direction, a pitch larger than half of the wavelength 133 can be chosen between the pMUTs of each channel. This 134 helps to have a sharper beam pattern in the elevation angle. 135 The number of pMUTs in each channel is also an important 136 factor in enhancing the output pressure and the elevation 137 directivity [18]. However, if the size of the transducer in the 138 elevation direction is too large, the lateral spatial resolution 139 will be negatively affected. Depending on the application 140 and depth of penetration, the elevation size of the transducer 141 should be chosen. Also, a larger number of pMUT cells in 142 a channel implies larger parasitic capacitance which leads to 143 higher power consumption and lower receive sensitivity and 144 signal to noise ratio (SNR).

145
The proposed high frequency pMUT array consists of 146 128 channels with a pitch of 200 µm. Each channel comprises 147 several pMUTs that are divided into three groups (aligned 148 with the elevation direction), in order to have access to each 149 group separately. The pMUTs in each group are connected 150 to each other and are accessible via a bond-pad. In this way, 151 the beam pattern in the elevation direction can be steered and 152 focused up to a certain level and make the array independent 153 of an acoustic lens. Fig. 1 shows a microscopic image of the 154 array, in which the inset shows a magnified view of some 155 of the pMUTs. Each pMUT has a diameter of 85 µm and 156 a Si membrane thickness of 6 µm. For low frequency pMUT 157 array with 128 channels with a pitch of 214 µm. Each pMUT 158 cell has a membrane diameter of 160 µm making this a 159 high fill factor design. Each channel consists of a column of 160 several pMUT elements connected to each other in line with 161 the z-axis.

162
A further difference between the design of the two arrays 163 is the presence of a thin polyimide layer in the high frequency 164 VOLUME 2, 2022  respective thickness of 30, 170, 30, and 110 nm was deposited 197 and lifted-off as the top electrode. The radius of the top 198 electrode was chosen to be about 70% of the membrane radius 199 to maximize the induced lateral stress and thus maximize the 200 displacement response [20]. Afterward, (e) a 1 µm polyimide 201 layer was deposited in order to (i) introduce more damping on 202 the pMUT membrane to widen the bandwidth and (ii) serve 203 as a protective layer since the array will be submerged in 204 ionized water or ultrasound gel. The polyimide layer is then 205 patterned to have access to the bottom and top electrodes, 206 which is required for wire bonding at a later stage. Then, 207 (f) the membrane was realized by a deep reactive ion etching 208 (DRIE) process on the backside of the wafer. This is followed 209 by HF vapor phase etching of the buried oxide (BOX) layer. 210 The cross-section SEM images of the realized membranes are 211 shown in Fig. 3. 212 Finally, in order to have electrical access to channels, the 213 transducer was wire-bonded to a PCB, where each channel 214 is connected to a coaxial cable. A part of the transducer wire 215 bonded to the PCB is shown in Fig. 4. The wire bonding is 216 later covered with an epoxy layer (EPO-TEK H54) for the 217 protection against water, humidity, and physical contact. The measured capacitance of each individual channel is about 221 8.5 nF for the low frequency design. However, for the high 222 frequency array design, in which the low pass filtering effect 223 is more prominent, we could successfully reduce the parasitic 224 capacitance from 8.5 nF to 1.8 nF per channel due to the 225     The transducer was also characterized underwater, as water 256 can be considered as a similar medium to a human body, 257 which gives a good indication of its performance for medical 258 imaging applications.

259
Firstly, transmitted acoustic pressure by the transduc-260 ers was measured. Fig. 6 shows a graphical illustration 261 of the underwater experiment. A commercial 1mm needle 262 hydrophone (from Precision Acoustics, UK) and a single 263 element transducer (PVDF, 19 mm diameter with 10MHz 264 VOLUME 2, 2022   The frequency response of the low frequency transducer is 288 shown in Fig. 7b, which was measured by sweeping the fre- and the center frequency, the impedance of the transducer was 298 estimated to be 1/2πf C = 8 . Table 1. shows the comparison 299 between specifications of the fabricated low frequency array 300 and other previously reported pMUT arrays in the literature. 301 It was tried to select the pMUT arrays that were intended to 302 be used in medical ultrasonic imaging applications.

303
There are three reasons why the high frequency array 304 has higher transmission sensitivity with respect to the low 305 frequency array. First, the high frequency array has a higher 306 input impedance. Therefore, since our signal generator has 307 an impedance of 50 , more voltage is transferred over 308 the pMUT. Second, the high working frequency, results in 309 a higher velocity of the membrane, which causes a higher 310 output pressure. And third, the low frequency array has a very 311 wide bandwidth, which means that its vibration is damped 312 significantly. Therefore, a lower vibration amplitude, and 313 consequently, a lower output pressure is obtained.

314
To measure the receive sensitivity of the pMUT transducer, 315 a single element wide-band transducer (Precision Acoustic, 316 UK) was used to transmit a burst signal. In this experiment, 317 an acoustic pressure wave was transmitted by the single 318 element transducer and received by one channel of the pMUT 319 transducer, which was connected to an oscilloscope termi-320 nated at high impedance. The transmitted signal was mea-321 sured by both the hydrophone and the pMUT transducer.  in Fig. 9b. Unfortunately, the pulse response of the single 339 element transducer, utilized in this experiment, is not short 340 and has some ringing. However, a very similar response was 341 captured by the pMUT channel. The reflection shown in the 342 received signal by the pMUT channel corresponds to the 343 reverberations bouncing back and forth between the pMUT 344 array and the single element transducer.  closer and two rods further from the array with a diameter of 380 7.5 mm and 5 mm, respectively. The imaging setup is shown 381 in Fig. 10b. A one cycle rectangular pulse at 1.5 MHz with 382 25 Vp-p was used to actuate each channel in the array. The 383 final reconstructed B-mode image is shown in Fig. 11b.

385
High frequency (6MHz) and low frequency (1.5 MHz) pMUT 386 arrays with 128 channels for medical imaging applications 387 are proposed in this paper. The design, fabrication process, 388 and characterization of the array are described in detail. 389 The pMUT arrays are characterized in air by LDV and 390 underwater by using a 1 mm needle hydrophone and a wide 391 bandwidth single element transducer. The functionality of 392 the arrays for imaging is confirmed by an underwater imag-393 ing experiment, which shows the high potential of the pro-394 posed pMUT array for medical imaging applications. Using 395 a high frequency array is useful to obtain higher spatial reso-396 lution images, while the energy of a lower working frequency 397 ultrasound wave is absorbed less by the medium and can 398 penetrate deeper into the body. The design of the arrays was 399 performed with an emphasis on the frequency bandwidth of 400 the pMUTs. A higher bandwidth in the frequency response 401 results in a shorter acoustic pulse and less ringing on the 402 output of pMUTs, which increases the axial spatial resolu-403 tion. A wide bandwidth is also necessary during the receive, 404 as it helps to capture other frequencies and harmonics gen-405 erated by the organs and tissues, which are useful for e.g.,

MICHAEL KRAFT leads the ESAT Research 549
Division Micro-and Nano-Systems. He joined 550 ESAT as a Full Professor in October 2017 and has 551 more than 20 years of experience in the design, 552 fabrication, and characterization of a wide range 553 of micro-and nanosystems (MEMS), sensors and 554 devices. He has worked on inertial sensors, intel-555 ligent interface circuits, and control systems for 556 micro-devices, atom and ion chips, bio-medical 557 and biochemical sensors and devices, energy har-558 vesters, and piezoelectric ultrasound transducers. He is in charge of the 559 Cleanroom and MEMS activities in the Leuven Nanocentre.