High Frequency Piezo-Composite Micromachined Ultrasound Transducer Array Technology for Biomedical Imaging

By Xiaoning Jiang

In this monograph, the authors reports the current advancement in high frequency piezoelectric crystal micromachined ultrasound transducers and arrays and their biomedical applications. Piezoelectric ultrasound transducers operating at high frequencies (> 20 MHz) are of increasing demand in recent years for medical imaging and biological particle manipulation involved therapy. The performances of transducers greatly rely on the properties of the piezoelectric materials and transduction structures, including piezoelectric coefficient (d), electromechanical coupling coefficient (k), dielectric permittivity (e) and acoustic impedance (Z). Piezo-composite structures are preferred because of their relatively high electromechanical coupling coefficient and low acoustic impedance. A number of piezo-composite techniques have been developed, namely “dice and fill,” “tape-casting,” “stack and bond,” “interdigital phase bonding,” “laser micromachining” and “micro-molding”. However, these techniques are either difficult to achieve fine features or not suitable for manufacturing of high frequency ultrasound transducers (> 20 MHz). The piezo-composite micromachined ultrasound transducers (PC-MUT) technique discovered over the last 10 years or so has demonstrated high performance high frequency piezo-composite ultrasound transducers.In this monograph, piezoelectric materials used for high frequency transducers is introduced first. Next, the benefits and theory of piezo composites is presented, followed by the design criteria and fabrication methods. Biomedical applications using piezo composites micromachined ultrasound transducers (PC-MUT) and arrays will also be reported, in comparison with other ultrasound transducer techniques. The final part of this monograph describes challenges and future perspectives of this technique for biomedical applications.

• Language: English
• Publisher: ASME
• Release date: Jan 1, 2017
• ISBN: 9780791861639

Book preview

1. Introduction and scope

1.1 Medical ultrasound

Medical ultrasound has been one of the most widely adopted and rapidly developing diagnosis and therapy tools today because of its non-destructive and non-ion radiative nature. Ultrasound transducers, as key components in medical ultrasound systems, have been developed for a variety of ultrasound 2D and 3D imaging with sub-millimeter to mm spatial resolution and Doppler blood flow information [1–4]. Medical ultrasound imaging stems from the sonar technology, which was first developed in World War I for detecting submarines using acoustic waves with frequencies ranging between >20 kHz and 1 MHz [5]. For a sonar transducer, a piece of piezoelectric material was used to produce and detect the acoustic pulses. The concept of sonar was later applied to image the human tissues by increasing the ultrasound frequencies above 1 MHz. Medical imaging with ultrasound was first demonstrated to be a useful clinical tool in the early 1950s by Wild and Reid, who reported the first two-dimensional imaging of soft tissues [6]. With the first medical ultrasound system, a transducer was scanned mechanically across the body to form a two-dimensional (2-D) image slowly. Later in 1970, the development of linear arrays and electronic beamforming made real-time 2D imaging possible [7]. The development of digital electronic beamforming in the 1980’s resulted in improved image quality and flexibility in scanning [8]. Ultrasound is now applied in a broad range of topics of biology and medicine, and accounts for about one-third of all diagnostic imaging procedures [8, 9].

1.1.1 Ultrasound wave basics and imaging modes

Ultrasound imaging usually involves ultrasound pulses, and the imaging resolution is mainly determined by the pulse duration (Figure 1-1). The wavelength λ of the propagating acoustic waves in a medium can be expressed as:

(1.1)

where v is the speed of sound in the medium and f is the ultrasound frequency. The ultrasound echoes received by the transducer are mainly the specular reflections from the medium interfaces [8]. The time that echoes arrives back to the transducer is determined by the distance between the transducer surface and the reflecting surface and the speed of sound in the medium, which is approximately 1500 m/s for tissue. The brightness corresponding to the amplitude of the echo depends on the mismatch of the acoustic impedance at the boundary and the distance between the interface and the transducer. In this way, time-to-distance and amplitude-to-boundary correlations are then established to express in the imaging information, in which an image line is noted as an A-line, where A is noted as Amplitude.

Figure 1-1 Diagram of the transmitting wave and receiving echoes. (Top) An ultrasound transducer is excited with a voltage pulse, and transmits an acoustic pulse into the tissue. (Bottom) The pulses reflected from the tissue are detected by the transducer corresponding different targets.

By acquring A-lines from a series of scanning positions and stacking them together, a 2D plane image can be formed. The brightness of the image is used to show the contrast, known as brightness mode, or B-mode. B-mode refers to the image plane that is perpendicular to the transducer surface, and C-mode displays the 2D image of plane parallel to the transducer surface. Each data point of the C-mode image is extracted from a A-line and corresponds to the certain depth of the interested plane, and the tranducer scans over the entire region to form the 2D view. Apart from the static image reconstruction introduced above, the ultrasound is enabled to show the motion information of the target, known as the M-mode, or motion mode. At the given position, the transducer emits and receives the pulse in quick succession, to capture the imaging lines with a short time interval. Using M-mode in echocardiography displays the movement of the myocardium, thus allowing for accurate and real-time measurements of wall thickness, internal diameter, and heart rate.

Doppler effect in medical ultrasound is greatly leveraged by doctors in the blood flow assessment. By determining the change of center frequency between emitting waves and receiving echoes, flow speed and direction can be estimated. One major application of doppler ultrasound is the detection and measurement of decreased or obstructed blood flow in vessels. Color Doppler ultrasound is done first to evaluate vessels rapidly for abnormalities and to guide the placement of the pulsed Doppler for detailed analysis of blood velocities.

1.1.2 Imaging resolution

The quality of the ultrasound image greatly relies on the resolution. The axial resolution is the closest distance between two objects that can be distinguished along an image line perpendicular to the transducer surface. The axial resolution is determined by the time duration of the pulse-echo response of the transducer [8]. The pulse length T is the temporal width of the envelope at half the maximum (–6 dB width). Two point targets will appear as individual reflectors in an image if the reflected signals are separated by a time greater than T. The corresponding spatial axial resolution Raxial is given by the following equation:

(1.2)

where v is the speed of sound in the medium. A shorter pulse provides a better axial resolution. The limit of the axial resolution is half of the wavelength in the medium. Usually, a typical pulse has 2–3 wave cycles.

The lateral resolution of a transducer is the minimum distance between two targets which can be differentiated on the plane parallel to the transducer surface. The lateral resolution RL is given by the following equation for a focused transducer [8]:

(1.3)

where λ is the wavelength in the medium and the f-number is the ratio of the focal distance r to the transducer aperture D. The lateral resolution can be improved by reducing the focal distance of the transducer, increasing the aperture of the array, and increasing the operating frequency. For a fixed number of wave cycles in each pulse, frequency increase would result in a decreased pulse length and narrower lateral beam, and hence the lateral imaging resolution, can be enhanced. For instance, As ultrasound frequency is increased to 50 MHz level, an axial resolution and lateral resolution of better than 20 and 100 mm for an f-number of 2.9 can be achieved [10], which are about 10 times smaller than those of a 5 MHz transducer with the same f-number.

1.2 High-frequency ultrasound

In recent years, high frequency ultrasound (>20 MHz) has been extensively studied and used in medical imaging including ophthalmic imaging [11], dermatology [12], catheter-based intravascular evaluation [13], and small animal imaging [14]. Intravascular imaging with probes mounted on catheter tips at frequencies ranging from 20 MHz to 60 MHz has been used to characterize atherosclerotic plaque and to guide stent placement and angioplastic procedures [15]. The medical efficacy of ultrasonic imaging has also been demonstrated in the anterior segments of the eye at frequencies higher than 50 MHz in diagnosing glaucoma, ocular tumors, and assisting refractive surgery [16]. The availability of a noninvasive imaging tool for dermatological applications could reduce the number of biopsies that are associated with patient discomfort and could better demarcate tumor involvement [12]. Small animal imaging is another major application of high-frequency ultrasound. Small animal imaging has been of intense interest recently because of the utilization of such small animals as mice and zebrafish in imaging, drug and gene therapy research [17].

Figure 1-2 Diagram of the enveloped receiving signal and imaging formation. (Top) The time and amplitude of the detected signals which represent the position of the reflector and strength of the echo. (Bottom) The signal processing from the enveloped amplitude to the brightness of imaging lines.

For high-frequency ultrasound imaging using piezoelectric transducers, the performance of transducers is greatly influenced by the properties of