Transducers used in the Cardiac Ultrasound Machine.
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Transducers used in the Cardiac Ultrasound Machine.
Abstract: Ultrasound imaging depends on the ability of piezoelectric crystals to generate sound when excited with alternating current and the reverse effect of charge accumulation or current flow when such crystals are subjected to pressure from sound waves. The first known ultrasound imaging machine was designed by K. T. Dussik in Australia in 1937. However, despite its widespread acceptance today, medical ultrasound did not develop as rapidly as X-ray imaging. Despite the relatively slow start, medical ultrasound imaging is very widely accepted today because there is no ionising radiation involved and hence the procedure is relatively safe. Ultrasound equipment is also cheaper as compared to X-ray imaging, magnetic resonance imaging, MRI and other techniques associated with nuclear medicine. The procedure involves minimal patient discomfort and is very useful for examining the soft tissues or the developing foetus. A dramatic increase in the number of older patients with chronic heart and valve disease has resulted in a prolific demand for the ultrasound cardiac imaging machines which can satisfy the requirements associated with fast and cost effective measurement of cardiac anatomy or function. One of the critical elements in the medical ultrasound imaging system is the ultrasound transducer without which signal processing and visualisation of the soft tissue images is impossible. Although many naturally occurring substances such as quartz exhibit the piezoelectric effect, lead zirconate titanate (PZT) ceramic ferroelectric materials have for many years been used for biomedical applications because of their superior characteristics for soft tissue imaging. Polyvinylidene difluoride (PVDF), transducer material has demonstrated advantages as a high frequency receiver. Single or multilayer transducers made of these elements can be used for ultrasound imaging as single transducers operating in A-mode or a two or three dimensional transducer array for B-mode, C-mode or M-mode ultrasound imaging. This brief essay takes a look at transducers for medical ultrasound.
The principle of operation of a cardiac ultrasound imaging device is based on the information that is provided by the varying delay times of echoes that are reflected from various depths of the human body tissue as a result of the ultrasound pulses that are generated by an ultrasound transducer being incident on the body tissue. Delay times of echoes from different depths are different and ultrasound is reflected from the interface of different types of tissues. A Doppler shift in frequency is also generated as a result of moving objects and the attenuation of ultrasound waves depends on the type of tissue that the ultrasound wave is travelling through. The ultrasound transducer which is responsible for the generation and detection of reflected ultrasound is, therefore, an essential component of the ultrasound imaging device. Ultrasound transducers work on the basis of the piezoelectric effect in which an alternating voltage applied to piezoelectric crystal material causes the crystals to become electrically polarised as a result of the applied electric field and hence vibrate with the alternating voltage to produce sound. Such crystals also become electrically polarised when stress is applied to them and hence any sound waves which are incident on them result in charge accumulation on the crystal surface and hence the generation of an alternating voltage. Thus, an ultrasound transducer consists of a suitable piezoelectric material sandwiched between electrodes that are used to provide a fluctuating electric field when the transducer is required to generate ultrasound. When the transducer is required to detect ultrasound, the electrodes may be used to detect any fluctuating voltages produced as a result of the polarisation of the crystals of the piezoelectric material in response to incident sound which generates fluctuating mechanical stresses on the material. Piezoelectric materials include quartz, ferroelectric crystals such as tourmaline and Rochelle salt as well as the group of materials known as the piezoelectric ceramics, which include lead titanate (PbTiO3) and lead zirconate (PbZrO3). These materials are also known as piezoelectric ceramics which are used in ultrasound transducers for biomedical applications.Polyvinylidene difluoride (PVDF) is another transducer material which has demonstrated advantages as a high frequency receiver. Piezoelectric ceramics are sold with the brand name PXE by Philips Company and are solid solutions of lead titanate (PbTiO3), and lead zirconate (PbZrO3) which have been modified by additives which are a group of piezoelectric ceramics known as PZT. PXE materials are hard, chemically inert and unaffected by a humid environment.
The crystals in a ferroelectric material of which PXE is made up of align themselves randomly in a number of directions. With such a random orientation of crystals, the material will exhibit no piezoelectric effect. In order to have a piezoelectric material which is capable of being used for ultrasound transducers, the material has to be subjected to a strong electric field at high temperatures. This has the effect of permanently locking the crystals in the direction of the applied electric field and making the crystal piezoelectric in the direction of the electric field. Hence, a piezoelectric ceramic material may be converted into a piezoelectric material in any given direction by applying a strong electric field to the material in the given direction at an elevated temperature. This treatment, which is known as poling, is the final stage in the manufacture of a PXE piezoelectric. Metal electrodes perpendicular to the poling axis are deposited on the material so that an alternating electric field may be applied to generate ultrasound or ultrasound vibrations may be sensed by sensing the electric field across the piezoelectric material. The voltage across a piezoelectric ceramic PXE material is usually directly proportional to the applied stress. The construction of a simple, single element piezoelectric transducer is as shown below.
The Construction of a Single Element Piezoelectric Transducer
Ultrasound imaging in the A-mode directs a narrow beam of ultrasound into the tissue being scanned and the echo which may be displayed on a CRT screen provides a measure of the distance between reflecting surfaces in the body. In the B-scan mode, the echo signal is brightness modulated which makes it possible for information related to tissue depth to be displayed on the screen in a visually effective manner. An ultrasound transducer array operating in B-mode permits a picture of the tissues within a patient’s body to be displayed on a CRT device. M-mode ultrasound imaging presents tissue movement by scanning an A or B – line on a monitor as a function of time and movements in this line indicate movements in the tissues within the body. In C-mode ultrasound imaging a second transducer is used to detect echoes sent out by the first transducer, presenting a 2-D map of the ultrasound attenuation within tissues.
Having discussed the principles of operation of a piezoelectric medical ultrasound transducer, it is now appropriate to consider the practical problems associated with the construction of such transducers. This is done below.
The Design of Ultrasound Transducers
A transducer which is constructed out of piezoelectric material will have a natural frequency of resonance and it is appropriate that the transducer should be excited with alternating electric field which matches the natural resonant frequency of oscillation of the material. The ultrasound frequencies that are used in medical imaging applications range from 1 MHz to 15 MHz and echocardiography is usually performed at frequencies of 2.5 MHz. Hence, transducers which are used for ultrasound imaging have to be tuned for different frequencies. For a transducer material in which ultrasound waves travel at the speed c, with a resonant frequency f, the thickness of the material is related by the formula f=c/2d. Hence, it is possible to tune various transducers constructed of the same material to different frequencies by adjusting the thickness of the material. The ultrasound transducer can be excited by a continuous wave, a pulsed wave, or a single voltage pulse depending on the requirements for observing a continuous image, echo ranging or other tissue measurements. The rear face of the piezoelectric crystal material is usually supported by a backing material which is tungsten loaded araldite, so that the vibrations in the piezoelectric material are rapidly damped after the initial excitation. It is important to couple the piezoelectric transducer to the body of a patient so that the incident ultrasound energy can be effectively transmitted into the body tissue that is being scanned. In order to do this, matching layers of suitable acoustic material are used along with a gel which makes it possible for the ultrasound waves to penetrate the tissue more efficiently. As far as possible, the characteristic acoustic impedance of the tissue being scanned is matched with the acoustic impedance of the transducer. The characteristic acoustic impedance of the tissue is defined as:
In the formula, c is the speed of ultrasound in human tissue which is about 1540 m/sec with a variation of +/- 6% and is the tissue density. K is the bulk elastic modulus of the tissue being scanned.
The acoustic parameters of an ultrasound transducer include its nominal frequency, the peak frequency which is the highest frequency response measured from the frequency spectrum, the bandwidth of the transducer which is the difference between the highest and the lowest – 6 dB level in the frequency spectrum, the pulse width response time of the transducer, which is the time duration of the time domain envelope which is 20 dB above the rising and decaying cycles of a transducer response, the loop sensitivity for a medium on which a test is performed which is characterised by:
Here, Vo is the excitation pulse voltage in volts, while Vx is the received signal voltage from the transducer. The signal to noise ratio for a biomedical ultrasound transducer is also an important parameter for an ultrasound transducer and this is defined as:
In the above expression, Vx is the received signal voltage from the transducer in volts in response to a specified tone burst or pulse and Vn is the noise floor in volts. The signal to noise ratio for an ultrasound transducer is a measure of the noise associated with the transducer, measuring instrument or cables and this is a good measure of how sensitive a transducer is. In addition to the previously mentioned parameters, geometrical parameters for a transducer describe how the acoustic pressure generated by a transducer varies across the axial and cross-sectional fields of a transducer. These variations are illustrated below:
Axial Beam Profile for an Ultrasound Transducer
Cross – Sectional Beam Profile for an Ultrasound Transducer
he detailed construction of an ultrasound transducer for medical applications involving the shaping of the piezoelectric material, matching layers, housing and backing materials etc is presently conducted using computational techniques such as Finite Element Modelling of ultrasound transducers through the use of software packages such as Ultrasim and other commercially available software. In the overall design, efforts have to be made to ensure that the overall design will be optimised so as to deliver a sufficiently high power of ultrasound into the tissue being imaged and as far as possible there is best possible sound impedance matching between the transducer and the scanned tissue. Design of the backing material in an ultrasound transducer is important because this design determines the ring down time of the transducer, which is critical for low noise and optimal axial resolution of the transducer.
Trends in Transducer Design for Echocardiography
Only the simplest equipment for echocardiography will use a single ultrasound transducer and there is a trend towards design of echocardiography equipment which uses two or even three dimensional arrays of ultrasound transducers to provide superior quality 2 –D or 3-D computer generated pictures of the organ being imaged. Even the relatively simpler equipment being used these days has two or more ultrasound transducers fitted into the transducer probe. The array of transducers are capable of generating a shaped beam of ultrasound which can be appropriately focused using electronic digital signal processing techniques to provide better images and resolution. Although the relatively simple medical ultrasound scanners cost about £1000 per piece, reasonably decent transducer assemblies for a decent Philips or Toshiba ultrasound machines can cost £1500 for the transducer alone. Transducer arrays for two or three dimensional ultrasound imaging equipment can be much more expensive because of the large number of transducers that are employed in such imaging equipment.
For better quality ultrasonic imaging to be possible, there is a requirement for enhanced bandwidth transducers, higher frequency transducer arrays and sophisticated digital signal processing circuits. There is also a trend towards transducer miniaturisation which will make intracavitary, intraurethral, or intravascular ultrasound (IVUS) investigation possible. The current imaging frequency range of 1 MHz to 15 MHz is expected to be increased to 20 MHz to 100 MHz and at these frequencies, microsonography devices using miniature ultrasound transducers with higher sensitivities are expected to provide much better and higher resolution images using catheter based transducers which are less then 2mm in diameter and are capable of being placed in veins. New ultrasound transducer materials are likely to provide transducers which are far more sensitive then those available presently and consume lower power. These transducers can be operated from battery powered portable equipment and there are indications in literature that with the availability of such devices, it is likely that the stethoscope will be replaced by miniature ultrasound equipment. New trends in ultrasound transducer construction are also moving towards composite transducer construction in which a composite of two piezoelectric materials is used to design the transducer.
Ultrasound transducers are fairly rugged and the piezoelectric material does not loose its properties unless exposed to high temperatures approaching the Curie temperature for the material are reached or there are strong alternating or direct electrical fields opposing the direction of poling for the material. Mechanical stresses imposed on the piezoelectric materials should not exceed the specified limits and although the specified limits vary for different types of materials, mechanical stress in excess of 2.5 MPa may be considered as likely to cause permanent damage. Ultrasound transducers are capable of being designed to operate in liquid mediums and the piezoelectric material does not react with water or gel.
Materials with piezoelectric properties such as lead titanate (PbTiO3) and lead zirconate (PbZrO3) lend themselves to being treated by poling to generate as well as detect ultrasound waves when subjected to alternating electric fields or mechanical stresses. Ultrasound transducers can be made out of these materials and these transducers can be designed for specified resonance frequencies for use in medical imaging. The detailed design of such transducers is an exciting and involving undertaking which is capable of being assisted by finite element simulations. Advances in transducer design involving the use of new materials, miniaturisation and the use of arrays of transducers promises to revolutionise medical imaging in the future by providing high resolution 3-D ultrasound images and the field is full of promise for device designers as well as computer engineers of the future.
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