Biaxial Driving of Piezoelectric Actuators for Ultrasound

ABSTRACT

Over the years, piezoelectric materials have been used in a wide variety of application on the health care field in the form of ablation and therapeutic treatments, medical imaging and lithotripsy, just to mention some of them. A more efficient operation of ultrasound transducers translates into less power consumption to obtain the desired effect. Commonly used ferroelectric actuators are driven by applying the electric field along the poling axis in order to maximize their mechanical response. The biaxial driving technique enhances the mechanical response of a piezoelectric actuator by dephasing two orthogonal electrical fields to the propagation and lateral electrodes. In this approach, several parameters are going to be tested to improve the acoustic efficiency response of transducers biaxially driven. These results will lead to evaluate the application of the biaxial driven technique on ultrasound imaging transducers.

Table of contents

ABSTRACT

1.- Introduction

2.- Background

2.1 Ultrasound

2.2 Ultrasound in medical applications

2.2.1 Diagnostic ultrasound

2.2.2 Therapeutic ultrasound

2.2.2.1 Physiotherapy

2.2.2.2 High intensity focused ultrasound

2.2.2.3 Lithotripsy

2.3 Acoustic propagation and acoustic wave

2.4 Characteristics of ultrasound

2.4.1 Frequency – Period – Wavelength

2.4.2 Acoustic impedance

2.4.3 Acoustic intensity

2.4.4 Attenuation and Absorption

2.4.5 Specular Reflection and refraction

2.4.5.1 Fluid-fluid interface

2.4.5.2 Fluid-solid interface

2.4.6 Scattering

2.5 Piezoelectric effect

2.6 Ferroelectric materials

2.7 Piezoelectric Ceramics

2.8 Single element ultrasound transducer

2.8.1 PVDF Transducers

2.8.2 Piezocomposite transducers

2.8.3 Capacitive micromachined ultrasound transducers (CMUT)

2.9 Material parameters

2.9.1 Piezoelectric charge constant

2.9.2 Permittivity

2.9.3 Electromechanical coupling factor

2.9.4 Piezoelectric voltage constant

2.9.5 Dielectric dissipation factor

2.10 Biaxial driving technique

3.- RESEARCH OBJECTIVES AND APPROACH

3.1 Research methodology

4.- CURRENT WORK

4.1 Biaxial transducer manufacturing process

4.2 Biaxial transducer efficiency characterization

4.3 Frequency sweep experiments

4.4 Power variation experiments

5.- PRELIMINARY RESULTS

5.1 Power efficiency vs. Electrical power

6.- WORK PLAN

7.- CONCLUSIONS

REFERENCES

1.- Introduction

Piezoelectric materials are used in a wide variety of medical application such as medical imaging, ablation and lithotripsy therapy treatments, to mention some of them. Piezoelectric materials for medical ultrasound transducers require high thickness mode electromechanical coupling coefficient (k_t), high dielectric constant (K^T₃), low acoustic impedance, low dielectric loss and high mechanical strength [1]. Some piezoelectric materials can improve its dielectric, piezoelectric and ferroelectric properties by using a doping material such as Gd, Al or Ce [2-4]. Also, the use of different geometries and arrays on ultrasound transducers can enhance the electrical and acoustic performance [5]. The efficiency of a piezoelectric actuator is determined by the amount of electrical power converted into acoustic power. Higher efficiency translates into less electrical power needed to achieve the desired acoustic power, leading to a reduced heating on the transducer and less power consumption by the system.

A method to increase this efficiency is using the biaxial driving technique [6]. The biaxial driving technique is based on the polarization rotation achieved by dephasing two orthogonal sinusoidal electrical fields along the x (lateral) and z (propagation) axis of a piezoelectric actuator to reduce the coercivity produced by the ferroelectric switching resulting in an enhanced mechanical response [6-9]. Results on [6] showed that the output acoustic power from a biaxial transducer depends on the phase difference of the applied electrical fields, which is different at every transducer and needs to be characterized for all of them.

The goal of this research work is to demonstrate the feasibility of driving piezoelectric actuators with the biaxial driving technique. For this purpose, twenty biaxial transducers were fabricated and characterized. Transducer were cut to have resonant frequencies as similar as possible in both orthogonal directions. The acoustic effective power of the biaxial transducers was measured with a radiation force method and compared with transducers driven by using a single electrical field.

2.- Background

2.1 Ultrasound

Sound is a phenomena that involucres the propagation of a mechanical wave through a medium. It can be propagated thought liquid, solid or gaseous elastics media [10]. Ultrasound refers to an acoustic wave whose frequency is above the range of audible frequencies for human beings, which is approximately 20 Hz to 20 kHz. In acoustics, waves are classified as follow:

1. Longitudinal waves, which consists in alternated compressions and rarefactions along the propagation direction in which the particles transmit the wave when moving back and forth from their equilibrium position parallel to the movement axis of the wave [11]. The propagation speed is determined primarily by the properties of the medium. Will be changes in the particle displacement, velocity, density, pressure and temperature associated with the wave propagation. [12].
2. Transversal waves, which consist in perpendicular vibrations along the direction of propagation. In contrast to the changes in particle volume that occur for longitudinal waves, no density change occurs for transverse waves. Transverse waves cannot be propagated in liquids [12].
3. Rayleigh waves consist of surface acoustic waves and are the result of the combination of the longitudinal and transverse waves. Each molecule runs an ellipse as the wave passes through it. This phenomena is produced when a longitudinal wave is tried to be generated on a surface which is in contact with air [10].

2.2 Ultrasound in medical applications

Ultrasound in the medical field can be found in a wide range of applications which can be divided into two major categories:  Diagnostic and therapeutic. The frequency used on any specific application is based on the considerations of sound absorption, penetration and, in the case of diagnostic, resolution [13].

2.2.1 Diagnostic ultrasound

The diagnostic medical sonography (sometimes called ultrasonography), is an ultrasound-based image technique used to visualize subcutaneous body structures as tendons, muscles, joints and internal organs [14]. Sonographers are typically used with a hand-held probe (transducer) which is placed and moved over the patient body. Water based gels are used as a coupling between the probe and the body. Usually, the frequency range for ultrasonography lies between 2 and 18 MHz and is very effective for imaging soft tissues [14]. There are four different modes for ultrasound imaging:

• The A-mode (amplitude modulation) is the most basic of the four. It is primary used to make internal distance measurements.
• In the B-mode, an array of transducers scans a plane through the body by moving the probe over the surface. The received echoes can be reconstructed to produce a 2-D image.
• The M-mode (motion) employs brightness modulation and is used to observe the movement of anatomic structures [13].
• The doppler mode, as its name states, uses the doppler effect to measure and visualize the blood flow. It can also be used to detect fetal heart rate [13].

Because of the low cost of this technology and being a non-invasive technique with a quick scan, benefits of portability, real time examination, and that do not produces radiation, it is well accepted by patients [15].

2.2.2 Therapeutic ultrasound

The versatility of ultrasound in medical therapies relies in the interaction with the cells and tissues. Ultrasound can induce effects not only though heating, but by the generation of ultrasonic cavitation (growth and collapse or motion of bubbles), gas body activation or mechanical stress [16]. The type of effects produced by the interaction of the tissues with ultrasound waves depends on several parameters, but the most important are the frequency, duty cycle and amplitude of the ultrasound wave [17]. Therapeutic applications can be divided in thermal applications, including physiotherapy and high intensity focused ultrasound, and non-thermal applications, like lithotripsy.

2.2.2.1 Physiotherapy

Commonly known as therapeutic ultrasound. Its objective is to warm tendons, muscles and other tissues by improving blood circulation to accelerate healing. It uses a hand-held probe which is placed on the surface of the tissue by using a water-based gel. The probe is moved on circles around the surface of the treating zone by a technician. The level of benefits are uncertain and the risk of getting burned is low if it is properly applied [16].

2.2.2.2 High intensity focused ultrasound

The high intensity focused ultrasound (HIFU), is a thermal ablation technique generally used for oncological treatments. A focused transducer generates an ultrasonic beam which is concentrated on a focal zone. The temperature on the focal zone raises until a lesion of a few mm in diameter is produced on the zone without damaging the surrounded tissues. This method is approved by the FDA to treat uterine fibroids [18], essential tremors [19], pain palliation of bone metastatic cancer [20] and to produce ablation of prostatic tissue [21].

2.2.2.3 Lithotripsy

Known as shock wave lithotripsy, this technique uses ultrasound waves to destroy targeted kidney or bladder stones at a slow rate. Its effects are based on two fundamental mechanisms, shock wave-related effects and cavitation phenomenon [22].

2.3 Acoustic propagation and acoustic wave

The physical nature of the medium can be modeled as stationary, with a spatially varying density

[35]. Typical piezoelectric polymers have a crystalline region with a dipole moment randomly oriented without any mechanical or electrical poling process with a net dipole moment equal to zero in this condition. This type of structure is called the

α

-phase PVDF film and has not a piezoelectric response [36]. By applying a mechanical stretching and electrical poling under a high electric field, crystalline regions inside the bulk PVDF film will align in the electric field direction. This type of structure is called the

β

-phase PVDF film and exhibits piezoelectricity [36]. One of its most important applications is broadband hydrophones [35].

2.8.2 Piezocomposite transducers

Conventional piezoelectric ceramics such as PZT, lead metaniobate and modified lead titanates are useful for making ultrasonic transducers used in medical imaging because of their high electromechanical coupling coefficients. Their principal limitation lies in their high acoustic impedance which makes coupling to tissue difficult. Piezoelectric polymers, such as polyvinylidene difluoride (PVDF), present a different set of material properties. Their low acoustic impedance simplifies efficient coupling, but their low electromechanical coupling, low dielectric constant and high dielectric losses present additional limitations when integrating these materials into medical imaging transducers [37]. Simply stated, a piezocomposite is a combination of a piezoelectric ceramic and a non-piezoelectric polymer to form a new piezoelectric material with the properties enhanced. In general, however, the term piezocomposite applies to any piezoelectric resulting from combining any piezoelectric polymer or ceramic with other non-piezoelectric material [38]. The advantage of such arrangement is that it combines the superior piezoelectric properties of the chosen ceramic with the lower acoustic impedance of the polymer [12]. This type of transducers are fabricated with common piezoelectric materials, but not in one piece. The material is selected and filled with a vibration absorbent material to isolate every cut on the piezoelectric [10]. In a two-phase system, e.g., ceramic and polymer, there are ten possible connectivities, of which only a few are of practical interest for transducer fabrication [12].

2.8.3 Capacitive micromachined ultrasound transducers (CMUT)

CMUTS are micromachined transducers formed from numerous unit cells electrically connected in parallel, with each cell consisting of a metalized membrane suspended above a heavily doped silicon substrate (top and bottom electrodes, respectively) [39]. Fig. 6 represents the CMUTs unit cell.

Fig. 6.Schematic cross section of a CMUT unit cell [39].

During operation, a dc voltage is applied between both electrodes. This dc bias voltage causes the membrane to deflect downward to a static operating point which determines the sensitivity, frequency response and total acoustic pressure [40]. If an alternating voltage is superimposed on the bias voltage the modulation of the electrostatic force results in the vibration of the membrane with subsequent generation of ultrasonography at the same frequency of the modulation [41]. Conversely, if the biased membrane is subjected to ultrasound waves, a current output is generated as a result of a capacitance changes due to membrane vibrations.

The advances in microfabrication technology have made possible to build capacitive ultrasound transducers competing with piezoelectric transducers. CMUTs offers advantages of improved bandwidth, ease of fabrication of large arrays with individual electrical connections, and integration with electronics [39]. CMUTs are fabricated using silicon micromachining methods with submicrometer accuracy and uniformity, which allow fabrication flexibility. Since silicon micromachining determines the shapes, sizes and spacing between neighboring elements, flexible shapes and multiple elements with transducer spacing as small as 3 µm can be fabricated [40].

Besides medical and underwater imaging, CMUTS have some potential applications that include air-coupled non-destructive evaluation, ultrasonic flow meters for narrow pipelines, microphones with RF detection and smart microfluidic channels [41].

2.9 Material parameters

2.9.1 Piezoelectric charge constant

The piezoelectric charge constant d_xy represents, for the direct piezoelectric effect, the polarization generated per mechanical stress (T) applied to a piezoelectric material and for the inverse piezoelectric effect, the mechanical strain (S) experienced by a piezoelectric material per unit of electric field applied. The subscript x represents the direction of polarization generated in the material or the direction of the applied strenght when the electrical field, E, is equal to zero. The subscript y indicates the direction of the applied stress or induced strain, respectively [32].

2.9.2 Permittivity

The permittivity or dielectric constant

εxyz

for a piezoelectric material is the dielectric displacement per unit electric field. Z is equal to T to represent the permittivity at constant stress, and equal to S to indicate the permittivity at constant strain. The subscripts x and y represents the direction of the dielectric displacement and the direction of the electrical field, respectively. [32].

2.9.3 Electromechanical coupling factor

The electromechanical coupling factor

kxy

is an indicator of the effectiveness with which a piezoelectric material converts electrical energy into mechanical energy, or convert mechanical energy into electrical energy (direct and indirect piezoelectric effect, respectively). The subscripts x and y denotes the direction along which the electrodes are applied and the direction along which the mechanical energy is applied or developed [32].

2.9.4 Piezoelectric voltage constant

The piezoelectric voltage constant

gxy

is the electric field generated by a piezoelectric material per unit of mechanical stress applied or the mechanical strain experienced by the material per unit of electric displacement applied. The subscripts x and y indicates the direction of the electrical field generated in the material and direction of the applied stress or induced strain [32].

2.9.5 Dielectric dissipation factor

The dielectric dissipation factor,

tan⁡δ

, express the parasitic loss that results by subjecting a material to alternating electrical fields. It is a measure of the electrical loss of the materials. When an electrical field, E, is applied to an ideal dielectric material, the resultant charging current find itself out phase by 90° with the applied electrical field. However, in a real ferroelectric material, the current also has a loss component in phase with th applied E and the net resultant current makes an angle

δ

with the ideal charging current. This loss current is a result of the dissipation of energy heat. This loss can be interpreted using the parameter

tan⁡δ

which is the ratio of loss current I’’ and charging current I’ [32].

Typical values for some of the above parameters, for some important materials, are listed on Table 1.

TABLE 1

Piezoelectric Materials Properties [23]