IEEE TRANSACTIONS ON SONICS AND ULTRASONICS, VOL. SU-31, NO. 4, JULY 1984 295

The Influence of Front-End Hardware on Digital Ultrasonic Imaging

MARK E. SCHAFER, STUDENT MEMBER, IEEE, AND PETER A. LEWIN

Invited Paper

Abstract-Digital ultrasonic imaging systems arc critically dependent upon transducers and "front-end" electronics, for the generation of insonifying pulses, the detection of reflected echoes, and the conversion of data to digital form. The impulse response, frequency and spatial characteristics of the transducer, and the electrical characteristics of the analog-t()-digital interface arc important factors in image reconstruction and enhancement, on a par with signal processing techniques. Selected theoretical and practical aspects of front-end hardware, and their influ­ence on image quality arc reviewed. The theoretical basis of transducer design is briefly reviewed, including methods for prediction of trans· duccr performance, acoustic field distribution, and element configura­tion; the limitations of each method are carefully discussed. Next, the design of the interface electronics is examined, demonstrating the requirements for sampling rate, signal-to-noise ratio, dynamic range, impedance matching, etc. The design and construction of a phased array transducer system is used to illustrate the practical aspects of transducer and interface design, including the trade-offs involved. Finally, future trends in ultrasonic transducer design arc discussed, including new piezoelectric materials (e.g., rare earth piczoceramics and piezoelectric polymers and composites), as well as the impact of new integrated circuit techniques.

I. INTRODUCTION

U L TRASONIC imaging for medical and nondestructive testing purposes has made significant progress since its

very beginnings over thirty years ago [ l]. Much of the recent increase in the use of diagnostic ultrasonic imaging can be traced to improvements in image quality. The new digital imaging systems coming into use hold the promise of further enhancing image quality. In addition, new image analysis techniques are a key element in the "ultimate goal" of medical ultrasonic imaging, which is tissue characterization.

The flexibility inherent in digital imaging systems enables the designer to use a wide range of algorithms to reconstruct, enhance, and analyze ultrasonic images. However, the useful­ness of these algorithms is limited by the quality of the original data. Critical to any digital ultrasonic imaging system is the "front-end": the transducer and associated electronics which link the digital system to the medium under examination. Although imaging capabilities can be augmented by digital processing techniques, the front-end hardware often sets the system perfom1ance limits. In particular, the signal-to-noise

ratio (SNR), bandwidth, and dynamic range of the raw digital data all have an influence on the utility of certain image pro­cessing approaches.

Manuscript received March 30, 1984; revised May 18, 1984. The authors arc with the Department of Electrical and Computer

Engineering, Drexel University, Philadelphia, PA 19104.

This paper reviews selected aspects of front-end hardware design, with an emphasis on those portions which are essential to digital ultrasonic imaging. The primary purpose is to ac­quaint the digital system designer with the basic operation of transducers and their associated electronics, examine some of the common imaging and system compatibility problems, and discuss the impact of front-end hardware limitations on digital imaging.

Although most of the practical examples to be given in this paper are taken from medical imaging system design, the basic system concepts are, in general, applicable to nondestructive testing (NDT) as well. The principal differences are due to the medium being imaged, more specifically its acoustic impedance. The acoustic impedance in diagnostic ultrasound is approxi­mately 1.5 Mrayls, close to that of water, while in most NDT applications, the impedances are on the order of 30 to 40 Mrayls. (The rayl, the unit of impedance in the SI system, is equal to the product of material density and sound velocity, and is expressed as kilograms per meter squared-seconds.) Consequently the major difference in front-end design is with respect to the acoustic matching (as discussed in Section II). The matching of the transducer to the interrogated medium is less critical in NDT applications, because the transducer acous­tic impedances are around 30 to 35 Mrayls. Another important difference between medical and NDT imaging is that the acous­tic source used in NDT may employ shear or surface waves to interrogate the media [2]. An excellent comparison of the design principles behind transducers for medical and NDT use is given in [3]. Apart from the comments relating to trans­ducer design, the contents of this paper will apply to NDT systems as well as medical ones. This paper does not, however, address the specific design considerations of extremely high frequency transducers used in acoustic microscopy ( 100-500 MHz).

The signal flow path of a typical front-end is given in Fig. 1; the system shown illustrates the minimum hardware required to form a single ultrasonic scan line. The transmitter produces a series of pulses or gated tone bursts, which excite the trans­ducer. The transducer converts electrical energy from the transmitter into an ultrasonic wave launched into the interro­gated medium, and converts the returned echoes back into an electrical signal. The transducer performance is particularly important because its spati<J and temporal characteristics have a predominant influence on the quality of the final image pro­duced. The receiver not only amplifies the returned signal, but is also responsible for analog signal processing, such as time-gain compensation (discussed later in Section III-B). The analog-to-

0018-9537/84/0700-0295$01.00 © 1984 IEEE

296 IEEE TRANSACTIONS ON SONICS AND ULTRASONICS, VOL. SU-31, NO. 4, JULY 1984

TRANSMITTER

TtSSUE fRi:INSDIJC~ fol

RECEIVER ANALOG-TO-DJC>l TAL CONUERTf R

Fig. 1. Block diagram of basic front-end system.

TO DJGITAl. SYSTEM

digital converter (ADC) interfaces the receiver to the digital processing system. The performance characteristics of these components affect imaging parameters such as spatial resolu­tion, depth of penetration, signal-to-noise ratio, bandwidth, and dynamic range, as discussed in the following sections.

In addition to the elements shown in Fig. I, a practical imag­ing system may also require timing and control electronics, together with some fonn of beam steering (deflection), either mechanical or electronic [4], [49]. Neither topic is discussed further, except for the case of beam deflection with phased arrays, which serves to illustrate the complexity of currently used medical ultrasonic imaging systems.

The following sections first examine selected transducer de­sign trade-offs in relation to digital interfacing, using the ex­ample of a disc transducer with front and back matching. This transducer configuration is chosen because its design and per­formance are relatively well documented, and it fonns the basis for other more complicated designs. Techniques for predicting and measuring transducer response are also briefly mentioned. Next, the design of the associated electronics is examined; in particular, the influence each subsystem has on overall imaging performance. Phased array systems are discussed, since they constitute the greatest design challenge, and are well suited for use with digital imaging techniques. Finally, possible future trends in front-end hardware, including both new transducer materials, and integrated circuit technologies, are examined, along with their possible impact on digital imaging hardware.

Before discussing transducers and front-end electronics in detail, it is appropriate to briefly review a few general principles which apply to medical ultrasonic imaging systems and provide the basis for many of the digital imaging requirements discussed later. Medical ultrasonic visualization systems are commonly used to image both major organ structures (i.e., abdominal, cardiological, and obstetrical), and "small parts" (i.e., eye, thy­roid, testicle, etc.). Typical diagnostic frequencies range from 1-5 MHz for the major organs, to 5-15 MHz for small parts scanning.

The most important parameter of an imaging system is its spatial resolution, both axial and lateral. Axial resolution depends upon the duration of the interrogating acoustic pulse, while lateral resolution is a function of the ultrasonic beam­width, and therefore of the focusing properties of the trans­ducer. For advanced ultrasonic imaging systems, the axial resolution is typically 0.5 /I., while the lateral resolution is on the order of 2'A, in the focal region p, is the wavelength of sound in the interrogated medium). Resolution is directly proportional to frequency (since higher frequencies have shorter wavelengths). However, because the attenuation of tissue increases with frequency, there is an immediate trade-off between resolution and signal-to-noise ratio (SNR). For a

given system sensitivity, this translates into a compromise be· tween resolution and depth of penetration.

Ultrasonic imaging further requires a wide dynamic range, since losses from tissue attenuation (1 dB/cm/MHz to a first approximation [4]) and reflection losses at tissue interfaces (20-30 dB) can produce 80-100 dB variations in received reflected echo amplitude. A I 00 dB system dynamic range is considered the minimum requirement for diagnostic applica­tions. It is worthwhile to note that the frequency dependent attenuation of tissue can introduce a spectral shift in the re­ceived signal. This effect may be compensated for using the appropriate signal processing technique to optimize the signal­to-noise ratio.

All the above considerations apply to digital imaging systems, especially if advanced image processing techniques are em­ployed. In those cases, the signal-to-noise, frequency band­width, and dynamic ranges criteria may be even more stringent, as discussed in the following.

II. TRANSDUCERS

From the above discussion, it is clear that the basic require­ments of the transducer are to transmit short acoustic pulses and receive them with high sensitivity: this insures adequate axial resolution, signal-to-noise ratio, and depth of penetration. Focusing, by either mechanical or electronic means, is used to narrow the ultrasonic beamwidth and maintain adequate lateral resolution. These parameters, sensitivity, bandwidth, resolu­tion, and signal-to-noise ratio, fonn the image quality criteria which influence transducer design.

Front-end design centers on the transducer, because once the transducer has been designed, the rest of the front-end is con­structed to be compatible with it and the digital processing system. Transducer design will not be examined in detail, in part because a complete examination of ultrasonic transducers is beyond the scope of this paper, and in part because a number of excellent papers on the subject have recently been published [ 5] , [ 6]. Instead, the basic structure of the transducer will be presented, and the requirements for adequate transducer performance from a digital systems integration standpoint will be examined. A single element disc transducer is used as an example; this type of transducer is used extensively for medical imaging in the 1-10 MHz range, with a typical bandwidth of 40-50 percent of center frequency. Multiple element trans­ducers are considered in Section IV.

To meet the above mentioned image quality criteria, ultra­sonic imaging transducers have evolved greatly over the past three decades [7], into the general fonn shown in Fig. 2; the design of this type of transducer was thoroughly examined by Hunt et al. [5] and by O'Donnell et al. [6]. In this config­uration, the piezoelectric ceramic disc is designed for half-wave resonance at the desired frequency [ 6] , in order to maximize sensitivity. Focusing is accomplished mechanically by curving the front face of the transducer, or using an acoustic lens. By itself, the ceramic disc element has a number of drawbacks as an imaging device, principally a strongly resonant response {high Q), internal reflections at the front and back faces of the ceramic, and a poor acoustic impedance match with tissue. The high Q results in a narrow bandwidth and the internal

SCHAFER AND LEWIN: INFLUENCE OF FRONT-END HARDWARE

TISSUE

MATCH!NG LAY[R -PIEZOELECTRIC

CERAMIC ELECTRODES

BACKlNG MATERIAL

Fig. 2. Basic piezoelectric transducer configuration.

refections cause excessive ringing in the transducer impulse response, both of which lead to a degradation of the quality of the final image. The impedance mismatch (the acoustic im­pedance of typical piezoelectric ceramic material is approxi­mately 30 Mrayls, versus 1.5 Mrayls for tissue) reduces the acoustic coupling efficiency and worsens the ringing problem.

The front matching layer and backing material are widely used to overcome these problems. The backing structure pro­vides a well-matched attenuating medium to couple to the piezoelectric ceramic disc and damp out the internal reflections. Backing layer design and composition are extensively treated in [8]-[10]. The front layer serves to match the impedances of the ceramic and tissue; the improved coupling to the load increases the efficiency and reduces the Q. The design of the front matching, to a first approximation, follows classical quar­ter wavelength coupling theory [11], using a front layer imped­ance which is the geometrical mean of the ceramic and tissue. However, the use of backing material and pulsed excitation influence the design of the matching layer; these points have been considered by a number of researchers. A more accurate matching procedure utilizing multiple matching sections was suggested in [ 17] : the impedances of the sections were chosen such that the resulting transmission line reflection coefficients exhibited maximally flat or Tchebycheff response over the desired frequency range.

The use of matching layers and backings presents it own problems. Although the damping provided by the backing does shorten the impulse response, it does so at the expense of sen­sitivity, since much of the acoustic energy is lost to attenua­tion. Matching layers, although they improve sensitivity, are effective only over a narrow frequency range, and their design is frequently impeded by the lack of suitable materials having the required acoustic impedances. Furthermore, adding match­ing layers complicates the assembly of the transducer, raising the production cost and increasing the possible deteriorating effect of the glue bonds between layers. The glue bond thick­ness should preferably be less than one twentieth of a wave­length (typically less than 3 µin the frequency range used in diagnostic ultrasound), in order to have a negligible effect [ 5] , (12].

While sensitivity, dynamic range and impulse response can be varied by the use of front matching layers and backing material, the transducer's focusing properties are controlled not only by front face curvature, but also the transducer diameter (aper­ture) and operating frequency. Strengthening the focusing (by increasing the curvature) improves the resolution in the focal zone but, for a given transducer diameter, it generally worsens the depth of field. Increasing the transducer aperture size

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increases its sensitivity, (since there is more area available for transmitting and receiving the ultrasonic energy). However, the depth of field shortens with increasing transducer aperture. In addition, manufacturing and anatomical constraints can limit the transducer dimensions used in practice. The trade-offs involved in the focusing of transducers, together with a com­parative analysis of some typical transducer geometries, are given in (5].

From the above, it is easily seen that there are trade-offs to be considered in the design of transducers. As already men­tioned, the goal is to produce a transducer which satisfies the requirements of the digital imaging system in terms of impulse response, sensitivity, signal-to-noise ratio, and beamwidth. The exact performance goals depend upon the digital techniques implemented, and have to be set by the digital system designer. Once the requirements are established, the transducer design trade-offs can be weighed against performance goals by the use of computer simulation, which is briefly discussed in the following.

The frequency response, sensitivity, and dynamic range trade­offs can easily be examined using equivalent circuit models. These models assume that the piezoelectric ceramic disc is thin and laterally clamped. It is further assumed that one-dimen­sional wave propagation dominates and that the involved struc­trues all act like acoustic transmission lines; this is generally the case for typical disc tranducers. The models most widely used are the Mason lumped parameter model [13], and the Krimholtz, Leeman, and Mattheai (KLM) transmission line model [ 14] . Both Hunt et al. [ 5] and O'Donnell et al. [ 6] give comparative analyses of these models and discuss their limitations. The models not only can simulate the piezoelec­tric ceramic material, but also can simulate the matching layers, backing structure, and the interrogated medium (tissue).

The transducer focusing properties can conveniently be found from an analysis of the ultrasonic beam patterns. The standard technique uses Huygens' principle, in which the transducer face is modeled by an infinite set of discrete point radiators, each generating a hemispherical wave. The contribu­tions from each radiator are summed in a diffraction integral to find the radiated acoustic field. This integration has been performed for a number of geometries, notably by Weyns [l 5a], [ l 5b]. Other extensive reviews of this approach can be found in (3], [16].

Various investigators have used these modeling techniques to produce transducers with broad-band low-loss characteristics (17] - [ 19]. The common design strategy, as outlined in [20] , is to use the lightest damping which still provides adequate impulse response, to achieve the best sensitivity. Matching layers are then used to further improve sensitivity. Concur­rently, previously mentioned assembly constraints, such as material availability and glue thickness, have to be taken into account. Other approaches to this design trade-off involve electronic circuits to improve the transducer response (21], (22], or variations of the excitation signal from a unipolar impulse to a bipolar pulse train (23] to reduce ringing ( de­scribed in Section III-A). These approaches generally result in improved sensitivity compared to heavily damped transducers, although complex electrical matching may not always be

298 IEEE TRANSACTIONS ON SONICS AND ULTRASONICS, VOL. SU-31, NO. 4, JULY 1984

possible for applications where space is restricted, such as multiple element tranducers.

In mechanically steered transducer imaging systems [24], [25], (phased arrays are discussed in Section IV), the design considerations are aggravated by the presence of the scanhead and the servomechanism. The scanhead includes the transducer, the coupling medium, and the contact (acoustic) window. These may often lead to reverberation, mode conversion, and beam distortion, all of which must be compensated for in the overall scanhead design.

Another important consideration for high quality transducers is concerned with radio-frequency (RF) shielding. This shield­ing is necessary to maintain the signal-to-noise ratio of the transducer, especially in the clinical environment, known for noisy electromagnetic interference conditions. The problem is especially noticeable in phased array transducers (discussed in Section IV), where the radio-frequency interference tends to produce ghost images along the scan line corresponding to zero-degree deflection (normal to the transducer face).

Before examining front-end electronics, it is worthwhile to briefly mention some of the techniques used to experimentally determine transducer beam patterns and impulse response. A widely accepted technique uses the transducer in pulse-echo mode, and places a reflector, such as a lucite or stainless steel plate or a stainless steel sphere, in the focused region or in the transition region for a nonfocused transducer. (The transition region is, to a first approximation, considered to take place at the axial distance a2 //...,where ais the largest transducer dimen­sion, and/... is the acoustic wavelength in the interrogated me­dium [4], [26] .) The received voltage output of the trans­ducer (either a waveform or a spectrum) is then recorded for various target positions, using an oscilloscope or a spectrum analyzer. Although relatively simple, the technique suffers from errors introduced by the reflection frequency response of the target. A technique which is gaining interest uses (minia­ture) piezoelectric polymer polyvinylidene fluoride (PVDF) hydrophone probes [27]-[29] to map out the transmitted field. The specific advantages of the probes include their wide bandwidth (typically 0.5-20 MHz), small size (0.6-1.0 mm diameter), and excellent directivity patterns. One immediate and very recent result of using the wide-band pro bes has been in acquiring evidence of significant nonlinear propagation ef­fects generated by standard ultrasonic imaging equipment [30], (31]. Nonlinear propagation leads to the transfer of energy from the fundamental wave frequency into its higher harmonics. Since medium attenuation is frequency dependent, this causes increased attenuation for the wave as a whole. As a result, the increased attenuation effectively reduces the overall sensitivity of the transducer, since nonlinear "excess absorption" can account for up to a thirty percent increase in total absorption (32]. These nonlinear effects were not noticed previously because of the limited bandwidth of the ceramic transducers probes available at that time.

A final transducer design consideration is associated with the limitation on transducer acoustical output power. Although increasing the transmit power increases the depth of penetra­tion of the imaging system, there are limits imposed by the aforementioned nonlinear nature of the medium, and concern

over possible bioeffects at high acoustic power levels. The acoustic output measurement methods and safety guidelines are extensively discussed in [33] and [34].

III. ASSOCIATED ELECTRONICS

In order to maximize the information transfer from the trans­ducer to the digital processing system, careful attention must be paid to the remaining parts of the front-end: the transmitter, the receiver, and the analog-to-digital converter. These three subsystems must match the performance characteristics of both the transducer and the digital system.

A. The Transmitter

The transmitter circuit has a number of design objectives which follow from those discussed earlier in connection with the transducer. Of these, in particular, efficient voltage excita­tion is of primary importance in order to maximize the SNR for a given transducer sensitivity. Typical transmitters fall into two basic categories: the pulse-type and burst-type. The former uses a single electrical spike or pulse in order to mini­mize the time duration of the transmitted ultrasonic pulse, and the latter generates a gated sinusoid or square wave in order to narrow the spectrum of the transmitted ultrasonic wave.

Pulse-type transmitters work by "shock exciting" the trans­ducer: a short high-voltage unipolar pulse is delivered to the transducer which resonates in its fundamental mode. Typically, these transmitters rapidly discharge a capacitor across the transducer using an electronic switch. The turn-on time is generally quite fast, on the order of 1-10 ns, depending upon the switching device used. The studies published include descriptions of transmitter switches employing silicon con­trolled rectifiers (35], an avalanche discharge scheme of bipolar transistors (36], and field effect transistors including MOSFET's (37]. A thorough comparative analysis of these switching devices [38], concludes that FET's, including HEXFET's and VMOS devices, are best for most ultrasonic applications, because of their fast switching time (I 0 ns) and low "on" resistance (approx. 5 ohms).

Since the excitation pulses are very narrow, and last for less than 100 ns, they have a wide frequency spectrum, generally wider than the transducer itself. Thus some of the excitation energy is wasted. The problem is compounded by the fact that the transducer is often well damped to reduce excess ringing, which in turn reduces efficiency. High excitation volt­ages (up to 400-1000 V) have been used to overcome these losses and increase the energy delivered in the transducer's frequency band. Pulse energy cannot increase without limit, however, since high pulse voltages represent a source of elec­trical interference, and a potential shock hazard to the patient and operator. Moreover, high pulse voltages can gradually depolarize the piezoelectric ceramic material, reducing the transducer sensitivity with time. Also, the wide frequency spectrum of the pulse tends to excite undesired modes of vibration in the transducer, (especially the transducer elements of linear arrays, cf. Section IV) which in turn degrades trans­ducer performance by distorting the frequency response.

Another approach for pulsed excitation (23] uses less heavily damped transducers (for higher sensitivity), and shortens the

SCHAFER AND LEWIN: INFLUENCE OF FRONT-END HARDWARE

transmitted acoustic pulse by "bipolar" electrical excitation. The study used a transducer operating in pulse-echo mode, and examined the received signal as a function of excitation pulse shape. The "optimum" received signal (maximum amplitude and minimum duration) resulted from using a sequence of volt­age pulses of alternating polarity, in order to counteract and minimize the transducer ringing. Since less acoustic energy was lost in the backing, the excitation voltage could be lowered to around 40 V, which is highly desirable in view of the above mentioned problems with high excitation voltages. A major drawback of this method, however, is that the pulses have to be experimentally determined for each transducer in order to achieve an optimal impulse response.

The second group of transmitters (burst-type) employ a number of cycles of a square wave or a sinusoid modulated by a Hanning-type window. Square-wave excitation is frequently used, since it requires only a simple switching circuit [23). One reason for using sine or square-wave bursts is to precisely control the frequency spectrum of the transmitted energy. This spectral control has proven useful in multiple frequency image processing techniques for speckle reduction [39]. Burst­type transmitters are generally used in phased array transducers, to improve the signal-to-noise ratio during coherent signal summation (cf. Section IV).

Another type of transmitter circuit which was recently sug­gested uses a digital-to-analog converter (DAC) to tailor the transmit burst under computer control [40]. In this way, any number of relatively complex voltage waveforms can be gen­erated to excite the transducer, under the limitations imposed by the conversion speed and bit resolution of the converter. The purpose is to adaptively modify the excitation voltage waveform in order to minimize the duration of the transmitted acoustic pulse, essentially by using an inverse of the transducer impulse response. This approach has been proposed in con­junction with digital imaging techniques, because it holds promise for improving axial resolution through inverse filter processing.

B. The Receiver

The primary purpose of the receiver is to amplify the signals from the transducer, generated by echoes returning from within the interrogated tissue. In doing so, it must not compromise the frequency or signal-to-noise characteristics of the trans­ducer. The wideband amplifiers used in receivers are described in many textbooks [41], and will not be discussed here. In­stead, we concentrate on some specific requirements of pulse­echo receivers which are important in digital imaging, including power matching of the transducer and receiver to maintain good sensitivity and dynamic range, and analog signal pro­cessing prior to analog-to-digital conversion.

First of all, matching the receiver to the transducer is indis­pensable in order to optimize the overall signal-to-noise ratio and sensitivity. Typical transducers for medical imaging have noise levels of approximately lµV, and voltage levels (on re­ceive) on the order of lOµV. The received echo signal depends upon the excitation voltage and the sensitivity of the trans· ducer (expressed as a two-way insertion loss, typically l 40 dB). In addition, it depends on the losses in the medium due to

299

tissue attenuation and reflection losses at tissue interfaces, as noted earlier. The maximum SNR is achieved when the trans­ducer and receiver impedances are matched, and both the elec­trical and mechanical resonances coincide. Although to a first approximation the transducer can be considered as a purely capacitive load or source, with a typical capacitance of between 1000 and 1600 pF (I 0 mm diameter disc transducer), the matching is complicated by the resonant nature of the trans­ducer's input impedance. Matching techniques have ranged from simple inductive tuning, primarily useful near the reso­nance [5], to complex broadband matching networks [21], which provide a wider frequency response with minimum distortion. These techniques are used to yield an effective impedance of 50 to 75 ohms, although in practice the actual impedance may range from 15 to 100 ohms.

In ultrasonic imaging applications in which a single transducer is used to both transmit and receive, the receiver is overloaded during the transmitted burst. This problem is usually overcome by a resistor in series with a diode clipper, as shown in Fig. 3. However, the series protecting resistor further complicates the electrical matching problem. On one hand, it must be large enough to maintain overload protection while transmitting, and on the other hand, it must be small enough to minimize loss and excess noise while receiving. The steps leading to an appropriate determination of Rare discussed in [42]. This paper also offers an alternative circuit, based on a lumped circuit equivalent of a tuned switching stub, which can improve the signal-to-noise ratio by approximately 6 dB.

From the point of view of this paper, the most important function of the receiver is analog signal processing. This pro­cessing is performed to complement the image processing taking place in the digital system. The analog processing is performed to reduce either the dynamic range or bandwidth requirements of the analog-to-digital converter. Since the ultrasonic signal is attenuated with propagation distance, and therefore with time, receivers employ time gain compensation (TGC) circuitry in order to achieve the best possible image quality, specifically by maintaining image uniformity. Otherwise the apparent bright­ness of a reflecting object would vary with position in the imaging field. TGC is under operator control, and it has been suggested that the digital processing system govern the TGC characteristics for maximum signal-to-noise ratio. TGC circuits which employ balanced differential amplifiers to achieve low noise and a high gain-bandwidth product are discussed in [ 43], [44]. Another form of processing often included in receivers is signal compression, which is necessary because the strength of reflected echoes varies over several orders of magnitude, and available gray-scale images using CRT screens can only present data with intensity variations over about 25 dB or less [4]. The compression is again operator adjustable, and can be varied from linear to logarithmic. Finally, receivers can include envelope or peak detection circuitry, quadrature sampling [45], and optional filters enabling differentiation of the signal envelope to emphasize tissue boundaries. These techniques are often used to reduce the bandwidth requirement of the analog­to-digital converter, as discussed in Section III-C.

The exact nature of the processing taking place in the receiver is critical to subsequent digital imaging techniques; it has been

300 IEEE TRANSACTIONS ON SONICS AND UL TRASONJCS, VOL. SU-3 J, NO. 4, JULY 1984

ro

HHIT(CllNG ~ESISTOR

Fig. 3. Standard circuit to prevent receiver overload.

pointed out [46] that acquiring infonnation on acoustical parameters of tissue in order to distinguish between normal and pathological states requires a detailed knowledge of the receiver signal processing.

C. Analog-to-Digital Converters

The analog-to-digital converter (ADC) is responsible for trans­ferring the signal infonnation obtained from the receiver into the digital processing circuitry of the ultrasonic visualization system. ADC specifications, including frequency bandwidth, resolution, and conversion speed are crucial to the quality of the final ultrasonic imaged produced. Unlike the transmitter and receiver, which are designed, constructed and optimized for operation with the transducer, the ADC is usually a com­mercially available off-the-shelf component. Since the basic design and application of ADC's are covered in a number of handbooks [ 4 7] , the present discussion is primarily confined to the ADC characteristics which are important for digital imaging, rather than on ADC circuit details.

Key parameters for the ADC to be considered are the conversion speed (sampling rate), resolution (number of bits). and accuracy. The conversion speed governs the frequency bandwidth of the digital system; the number of bits establishes the dynamic range and quantization process and therefore the contrast of the final image. Since these parameters cannot be optimized simultaneously, certain trade-offs are unavoidable. As an example, direct sampling of the signal from a 5 MHz transducer with a 40-percent bandwidth requires a sampling rate of at least 12 MHz, and preferably 15 MHz. In practice, sampling can be perfonned at twice the bandwidth of the transducer and still give sufficient accuracy, which still leaves a 5 MHz sampling rate requirement. These sampling rates are needed to avoid aliasing, which can introduce image artifacts and cause blurring. Since at least 40 dB of dynamic range is required to image tissue interfaces, an eight bit converter (which corresponds to 48 dB dynamic range) is only marginally acceptable. More sophisticated processing to determine blood velocity (Doppler), or to analyze the image texture of organ structure (tissue characterization), requires even higher dynamic range. Generally there is a direct trade-off between the conver­sion speed of the ADC and dynamic range achievable, and converters with sampling rates in the low megahertz range have typically only four to eight bits. Most of the commercially available converters exhibit either low speed and high resolu­tion or high speed and low resolution, which is of limited help from the acoustic imaging point of view. The above discussion clearly emphasizes the importance of the analog signal pro­cessing performed in the receiver section in order to reduce the speed and resolution requirements of the analog-to-digital converter.

Parallel or "flash" converters [47] have the highest speeds

(over 50 MHz at four-bit resolution), but their complexity grows rapidly with increasing resolution, e.g., a four-bit ADC requires only 15 comparators, while an eight-bit ADC demands 255 comparators. This affects the total cost since more com­plex designs require, in general, additional adjustment during the manufacturing process. As a result, standard flash con­verters are rarely used for greater than four-bit resolution. ln practice, the optimum performance from a speed and resolu­tion point of view Is obtained by using two-stage parallel con­verters. This type of converter achieves a reduction in circuit complexity at the expense of sampling speed and has speed/ resolution parameters better suited to digital ultrasonic imaging. A recently introduced two-stage converter, for example, has twelve-bit resolution at a 10 MHz sampling rate [48] .

Ultimately, the choice of ADC depends upon the speed and resolution requirements imposed by the processing techniques employed in the digital imaging system. It is also important to stress the interaction between the receiver system design and the analog-to-digital converter, i.e., the resolution requirement is less critical when signal compression is performed in the receiver. This compression, however, means that subtle tex­tural information has been irretrievably lost.

IV. A PHASED ARRAY FRONT ENO

Phased array front-end systems were chosen here as an ap­propriate example in order to address the recent advances in ultrasonic imaging. In the past few years, phased array imaging systems, although in general more complex and expensive than mechanical or linear scanning technologies, are gaining atten­tion because of their flexibility. Phased array systems are closely linked to new digital imaging developments, primarily because it is possible for the image processing system to control the beam pattern adaptively. When compared to mechanical scanners, phased array systems offer several advantages. They allow time adjusted (dynamic) focusing, in which the position of the focal zone is changed with time to extend the depth of field. The effective aperture size may also be varied as a func­tion of time (dynamic aperture and dynamic apodization), which allows improvements in beam resolution and sidelobe levels (59]. Further, they contain no moving transducer parts, which increases their operational reliability. The disadvantages, however, include increased complexity and cost of the frnnt­end system.

The general arangement of a phased array front end is shown in Fig. 4. The transducer is composed of a number of individ­ual elements, each of which is operated independently. The transmitted burst from each element is time delayed in a spe­cific pattern across the array: if the delay pattern increases linearly, the beam is deflected; if the delay pattern is adjusted to account for the acoustic transmission delay between each element and a focal point, the beam is focused [49]. Similarly, the signals received from each element may be delayed and summed coherently to steer and focus the receive beam. Thus, by changing the transmitting and receiving delays; a sector scan can be made. Additionally, the receiver focusing ability can be varied with time (dynamic focussing) to extend the depth of field. This is a distinct advantage over the single

SCHAFER AND LEWIN: INFLUENCE OF FRONT-END HARDWARE 301

1RANSMlTTERS

DELAYS

PRCAHPS DELPYS ~ECEIUEA5

~·~ ANALOG-rO-Ditil TAL CONVERTER

• • . . . " . . TO DIGITAL SYSTEM • • • • • •

~~ COHERENT SUMMTJON

Fig. 4. Block diagram of phased array front-end.

element transducer discussed earlier, especially in real-time cardiological and neonatal applications.

A comprehensive review of the general principles of phased array visualization systems from an image quality point of view can be found in [49]; the paper also provides an excellent background on the physics of beam steering, and the relation­ship between the transducer resolutions (axial and lateral), and the spatial sampling rate. The following discussion, then, is confined to selected current problems in transducer design, and the additional complexity of the electronics used for phased array imagers.

A. Phased Arr~ Transducers

Fig. 5 shows the construction of a typical phased array trans­ducer. A sandwich structure of the piezoelectric ceramic including backing and matching layers is separated into indi­vidual elements. Commercially available phased array trans­ducers contain typically from 32 to 64 individual elements. However, systems have been proposed which comprise up to 256 elements. The number of elements is determined by limitations on total aperture size, which in turn is governed by both anatomical constraints and manufacturing considerations. The center-to-center element spacing is typically a half wave­length or less (250µ for a 3 MHz array) in order to prevent grating lobes, which cause ghost images and reduce the signal­to-noise ratio of the image. This spacing requirement has been one of the most significant limitations in using phased arrays at higher (7-15 MHz) frequencies. In addition, the directivity pattern of a single element, expressed as the sensitivity roll-off with deflection angle, affects the imaging performance by reducing the SNR as the beam is deflected [ 50] . It is clear that phased array design involves additional parameters apart from those already discussed in Section II. The design considerations fall into three basic groups: the change in the characteristics of the piezoelectric and matching layers caused by the nondisc geometry, the assembly and dicing of the individual elements and, finally, the mechanical, electrical and acoustical cross­coupling between elements.

Informed and careful design requires exact modeling, em­ploying one of the two equivalent circuits mentioned earlier in Section II. However, the equivalent circuit techniques used in the design of disc transducers assume that the element is suf­ficiently wide to permit an essentially one-dimensional treat­ment. As it appears from Fig. 5, this condition no longer holds in the case of phased arrays, since the requirements for half­wavelength element thickness (see Section II), and small

MATCHING LAVE~S

PIEZOELECTRIC CERAMIC

Fig. 5. Phased array transducer configuration.

(less than half a wavelength) center-to-center spacing result in elements which are thicker than they are wide. This geometry introduces undesirable parasitic modes of vibration which are not accounted for in the standard transducer models. The ratio of width to thickness (W/T) severely affects the interaction of the fundamental and parasitic modes. This interaction both reduces the transducer sensitivity by wasting energy in the parasitic vibration modes, and degrades the single element directivity by distorting the shape of the radiating face.

This vibrational coupling can be partially accounted for by incorporating modified piezoelectric constants in the equiva­lent circuit models described earlier. The modified parameters can be derived by considering the coupling between the desired extensional mode and the unwanted lateral mode as a function of the width-to-thickness ratio [51]. In addition, the method described in [ 51] accounts for the variation in acoustic loading of the medium and the backing caused by the narrowness of the element. The results showed excellent agreement between theoretical predictions and experimental measurements for both impedance and insertion loss. The modified constants can also be derived from measuring impedance data [52]. This technique is based on the Mason equivalent circuit [ 13] , and the parameters of the model are varied until the model predictions match the actual measurements. The "effective" parameters are then used to predict the effect of the trans­ducer geometry on its performance.

Another powerful analysis technique which has gained atten­tion in recent years is based on the finite element method (FEM). For example, Sato et al. [ 53 J discuss the use of FEM in the analysis of vibration modes in single elements used in phased array transducers. Although well suited for analysis of advanced piezoelectric transducers, FEM analysis requires significantly more computation time that the equivalent circuit methods, especially for a three-dimensional simulation of the transducer structure. A technique has been proposed which can reduce the computation time by a factor of two, by in-

302 IEEE TRANSACTIONS ON SONICS AND ULTRASONICS, VOL. SU-31, NO. 4, JULY 1984

eluding piezoelectric effects as a perturbation of the basic mechanical properties of the transducer [54]. The perturba­tion technique has been shown to accurately predict resonance frequencies and mode coupling coefficients for a number of dif­ferent geometries. It is worthwhile to mention that the accuracy of the perturbation technique is limited only by the number of finite elements used, not the perturbation assumption.

Separating the transducer into elements (dicing) is the second design consideration in phased array transducers, and consti­tutes a major challenge in their construction. In a 3 MHz phased array transducer,' for example, the slots between the elements cannot exceed SO µ, in order to avoid grating lobes. These narrow slots are achieveci through a precision diamond sawing technique, similar to that used in the semiconductor industry. Not only is high precision required, dicing the trans­ducer may result in the possibility of damaging the electrodes, or partially depolarizing the piezoelectric ceramic material, if the temperature is not properly controlled.

The final aspect to be considered in designing advanced multielement transducers, especially those for phased arrays, is the acoustical, mechanical, and electrical cross-coupling be­tween elements. Cross-coupling, which affects the input impedances of the transducer elements and increases their effective radiating width, must preferably be kept well below 30 dB for adequate imaging [ S 1] . A major source of the coupling is due to electrical interactions between single ele­ments. This type of coupling is usually examined using a semi­empirical approach by estimating the electric field associated with each element [55). This approach can be used to establish the minimum cut (slot) depth for adequate isolation, as a func­tion of voltage excitation levels and the dielectric properties of the piezoelectric material. The other sources of coupling arise from surface and Lamb waves generated in the backing material and front matching layers, respectively. A recently developed theory of acoustic cross-coupling in shear-wave transducers, (which can be applied to longitudinal wave transducers as well) [56), demonstrated that cross-coupling will always exist to some degree, but that it can be minimized by 1) using a low­impedance high-loss backing to reduce the mechanical energy entering the backing and thereby decreasing the energy avail­able for coupling, and 2) making saw cuts which extend into the backing (to a depth of I .5 to 2 wavelengths) to increase element isolation. The isolation may be further enhanced by tilting the saw cuts as shown in Fig. 6(a). An experimental investigation thoroughly examining cross-coupling phenomena found a resonance effect involving the symmetry about a single element and its two nearest neighbors [57). The recommended solutions included 1) a differential dicing pattern, (Fig. 6(b )), to destroy the symmetry about each element, and 2) a metal foil placed over the face of the assembly to stiffen the entire structure. Although this foil does stiffen the assembly and constitutes an excellent RF shield, it can support Lamb wave propagation along its surface. Larson [58), in his com­prehensive review of phased array transducer design, points out that the foil should be as thin as possible (less than 25 µ) to minimize the L~b wave effects.

As the preceding discussion indicates, the design of phased array tranducers is extremely complex, mainly due to the

TRANSDUCER ELCtlENT

DACIONG 11ATERIAL

(a)

TRANSDUCER j ELF.:HEHT ~

(b) Fig. 6. Slotting techniques to reduce cross-coupling. (a) Slanted cuts

(after (56)). (b) Differential cuts (after [57)).

geometry employed. Just as the transducer for a phased array system is significantiy more complicated than single element disc transducers, the electronics used in phased arrays is more involved. The next section addresses the electronics of phased array systems, and stresses their influence on the requirements for the ADC converter.

B. The Electronics of Phased Arrays

The electronics of phased arrays in complicated by the need to properly steer and focus the transmitted and received acous­tic beams. For the transmitter, this involves exciting each phased array transducer element in the proper sequence; simi­larly, for the receiver, it requires properly delaying the signals received from each element.

The conventional approach for generating the desired trans­ducer excitation pattern employs a separate transmitter for each transducer element, either burst-type or pulse-type, al­though burst-type seem to prevail in recent designs. (The rea­son for this is that burst operation increases the signal-to-noise ratio by increasing the coherent energy available for imaging.) A typical pulser circuit suitable for array systems can be found in [59), while a more sophisticated burst-type transmitter circuit which uses field effect transistor switches (specifically MOSFET's) is described in [60]. Another interesting solution to phased array transmitter excitation has been published recently [44]. The transmitter circuit suggested contains digi­tal counters controlled by the master high-frequency clock bus, which runs at twice the transducer center operating frequency. Each counter is preset with the appropriate delay value to properly sequence the array firing pattern. This approach re­duces the circuit complexity needed to properly delay the transmit sequence, while providing a high degree of flexibility in the choice of delay values. Another approach to precisely

SCHAFER AND LEWIN: INFLUENCE OF FRONT-END HARDWARE

control the excitation of single elements employs tapped digital delay lines (shift registers) (61]. This technique uses a multi­plex switching network to connect each transmitter driver to the appropriate delay tap. Although conceptually straightfor­ward, the complexity of this type of circuit grows rapidly with the number of transmitting channels.

As already noted, optimization of the generated acoustic signal waveform requires appropriate matching of the trans­ducer load to the transmitter and receiver circuitry. In phased array transducers, the matching is additionally difficult due to the packaging constraints imposed by the small transducer size. Therefore simple inductive tuning is frequently used instead of the more involved electrical matching techniques described earlier in Section III.

The receiving system is also significantly more complicated in the case of a phased array, since the system must now be capable of focusing and deflecting the received beam. In addition, the relatively small area of the single elements in­creases their electrical impedance (typically 100 to 1 SO pF) and makes impedance matching more difficult. The individual elements also require preamplification because of their rela­tively low output levels in receiving mode. These problems can be overcome to a certain extent by using inductive tuning (similar, but not necessarily identical to the transmitter tuning), and a bipolar transistor preamplifier to better match imped­ances with the transducer (59].

As previously mentioned, one of the key features of phased array transducers is their ability to dynamically focus and de­flect the acoustic beam, thus enabling a sector of the interro­gated medium to be visualized. Dynamic focusing can be done in receive only, or can use a combination zone focus on trans­mit and receive; both techniques have been implemented in commercial scanners (49]. The focusing and steering abilities are especially dependent upon the techniques used to delay the received signals. The delay must be variable, in order to change the focus and deflection angle, and may range up to ten to twenty times the period of the transducer center fre­quency, depending upon the size of the array and the maxi­mum beam deflection angle. The received signals are typically delayed using discrete-time-increment delay lines, either analog or digital, although the resulting quantization of the delay time can cause grating lobes and reduce the signal-to-noise ratio [62], [63]. The time delay resolution generally must be on the order of one tenth of a period of the transducer center frequency, to maintain adequate image quality. Although somewhat bulky and expensive, analog delay lines have been used successfully to implement the necessary delays, providing S MHz bandwidth for delays in the range of0.005-15 µs [ 49]. Reference [49] also compares different analog and digital delay devices used for performing electronic focusing, with regard to their bandwidth, accuracy, size, and cost.

There are a number of different array beamforming ap­proaches which may be applied to phased array systems. The different techniques are used to reduce the cost or complexity of the scanners, and to reduce the delay hardware bandwidth requirements. For example, the bandwidth requirements can be relaxed by using a quadrature sampling technique [45]. The quadrature technique was used to reduce the bandwidth

303

requirements by a factor of three, while producing beam pat­terns equivalent to those obtained from full-bandwidth analog delay lines. Another technique for reducing the bandwidth requirements of the delay lines mixes each receiving channel with a phased local oscillator to provide the desired signal phasing, and then uses a single delay line to align the signal envelopes [ 64] . This technique has been incorporated in a commercial 64-element phased array imaging system [ 65] , where it allowed a significant reduction in circuit complexity while maintaining both adequate signal-to-noise ratio and dynamic range.

It is here appropriate to reiterate that the analog echo signals received fonn the basis for the final acoustic image. These analog signals are converted into digital ones as the digital sig­nals are better suited for advanced signal processing. The analog-to-digital converter must therefore not compromise the quality of the received signals, in terms of bandwidth, dynamic range, or signal-to-noise ratio. The analog-to-digital converter system for phased arrays varies according to the beamforming techniques used, but in general, the ADC must meet the same dynamic range, bandwith and conversion speed requirements discussed in Section III. Combined digital and analog beam­forming naturally involves a more complicated design, but the basic trade-offs discussed earlier still apply. In view of the fact that the cost of commercially available digital components is decreasing, and the components are undergoing miniaturiza­tion, it is likely that in the near future, techniques which integrate ADC's into phased array receiver systems will become more prevalent.

V. FUTURE TRENDS AND CONCLUSIONS

Apart from the gradual improvement of the systems and techniques discussed earlier, advances in front-end hardware will most likely come from two areas: new transducer materials and the further implementation of integrated circuit techniques.

One of the new transducer materials is a modified piezoelec· tric ceramic material, referred to as a "rare earth" piezoceramic. These ceramics are produced by doping lead titanate ceramic (PbTi0 3 ) with one of the rare-earth elements, such as samarium. The preparation of the ceramics, and their electromechanical properties as a function of composition are reviewed in [ 66], (67]. While these ceramics have approximately the same electromechanical coupling as the traditionally used lead zirconate titantate (PZT) ceramics, they exhibit only one tenth the planar coupling to lateral vibration modes (cf. Section IV). This allows greater freedom in the design of multiple element linear array transducers, since less energy is lost in the undesired modes. The modified ceramics have been incorpo­rated in a linear array transducer, and have permitted width-to­thickness ratios of greater than unity with no spurious vibra­tions and excellent impulse response, indicating a significant reduction in mode coupling [68].

Another transducer material which recently has gained sig­nificant attention is a piezoelectric-polymer polyvinylidene fluoride or PVF2. References (69] and (70] are two excellent reviews of PVDF properties, processing techniques, and appli­cations to medical imaging. PVDF has a lower electromechan­ical coupling and a higher loss tangent than PZT-type ceramics,

304 IEEE TRANSACTIONS ON SONICS AND ULTRASONICS, VOL. SU-31, NO. 4, JULY 1984

which reduce its efficiency when transmitting. However, it has two features which make it useful in medial ultrasonic imaging: its acoustic impedance is relatively close to that of tissue, and it is mechanically flexible, thus facilitating new transducer configurations such as variable mechanical focusing. A number of researchers have exploited the physical flexibility of PVDF to form transducers with toroidal and conical shapes [ 5], [71] . This configuration has shown excellent focusing over an ex­tended depth of field compared to a standard disc transducer of the same diameter. Because of their intrinsic broad-band properties, PVDF transducers have short impulse responses without matching layers. In addition, PVDF transducer arrays can be made using photolithography techniques instead of the dicing method discussed earlier in Section IV [72], [73]. The PVDF arrays have shown excellent impulse response and low interelement coupling (typically greater than 30 dB isolation), but relatively poor overall sensitivity and dynamic range (approx. 35 dB), principally due to the limited material thicknesses then commercially available. Despite these draw­backs, preliminary investigations indicate that PVDF, with further development, will be a viable transducer array material, for example in "small parts" imaging scanners, which do not require large depth of penetration. PVDF will be especially important in the design of high-frequency arrays (well above 5 MHz). At these higher frequencies, the photolithographic technique will allow closer element spacing than currently existing piezoelectric transducer technology.

Another promising type of piezoelectric material which has been suggested for ultrasonic imaging applications is the so­called "composite" material [74]. In the composite material, small spheres or rods of piezoceramic are incorporated in a matrix material. By varying the size and number to the piezo­ceramic inclusions, and the mechanical properties of the matrix, the electromechanical properties of the composite can be tai­lored to the intended application. The result is a new material with the transduction capabilites similar to that of the piezo· ceramic inclusions, but with lower acoustical impedance and better physical flexibility for easier focusing.

The above examples illustrate some of the possibilities for improved transducer designs using new materials. In addition, front-end perfonnance can be improved by implementing integrated circuit techniques. This issue will be briefly reviewed below.

Integrated circuit techniques can mainly influence front-end design in two ways: integration of the front-end electronics and the use of integrated circuit (IC) techniques in transducer construction. It is conceivable that, as hybrid IC technology expands, greater portions of the front-end electronics will be incorporated into custom designed !C's. These integrated circuits would not only have the advantage of small size, but also improved frequency response and lower noise pickup, because of markedly reduced stray capacitance. Some prelim­inary development work on integrated front-end design has been reported [75], as well as the use of surface acoustic wave devices for integrated analog signal processing [76]. Incorporating multiple analog-to-digital converters in complex (e.g., phased array) design would allow for parallel processing, which in turn would increase the overall imaging speed, and

possibly allow simultaneous scanning in several directions. A phased array system which makes use of paralled processing techniques has recently been discussed in [49].

The implementation of IC techniques in ultrasonic transducer design and assembly has so far focused on the development of two-dimensional arrays. which would permit three-dimensional beam steering and focusing. One such system has been de­scribed [77], including details of the transducer construction with a planar technique, the development of custom IC multi· pl ex switches, and the perfonnance of the final assembly. The system tested had sensitivity comparable to currently available PZT transducers, but with significantly reduced (500 kHz) bandwidth. A two-dimensional array can also be produced from a stack of linear arrays, reducing the assern bly problems asso­ciated with the planar technique and pennitting the inclusion of matching and backing layers to improve the bandwidth of the ultrasonic imaging transducer [78], [79].

In conclusion, the influence of front-end hardware on med­ical ultrasonic visualization systems was outlined, with an emphasis on those aspects of front-end design which are of primary importance in digital ultrasonic imaging. The purpose of this review was to provide the digital system designer with some insight into the trade-offs involved in the interfacing be­tween the piezoelectric transducer and the subsequent digital signal processing circuitry. The basic structure of piezoelectric ultrasonic transducers was examined, along with a brief review of the theories used to predict transducer performance. The design of the front-end electronics was discussed; in particular the requirements for signal-to-noise ratio, dynamic range, sampling rate, and electrical impedance matching were thor­oughly considered. A phased array imaging system was used to illustrate the complexity of current front-end hardware design. Finally, future trends in piezoelectric transducer materials (including rare earth piezoceramics, piezoelectric polymers, and composites) and associated integrated circuit techniques were briefly reviewed.

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Mark E. Schafer (S'83) was born in Pittsburgh, PA, on February 25, 1958. He received the S.B degree in electrical engineering from the Massachusetts Institute of Technology, Cam­bridge, in 1979 and the M.S. degree in acoustics from the Pennsylvania State University, Univer­sity Park, in 1982. He is now a Ph.D. candidate working under a Calhoun Foundation Fellow­ship at the Biomedical Engineering and Science Institute, Drexel University, in Philadelphia, PA.

In 1979 and 1980, he worked as an Associate Consultant in environmental noise control at Bolt, Beranek, and New­man, in Cambridge, MA. In 1982, he worked as a Reseach Engineer for the Applied Research Laboratory, State College, PA, in the field of underwater acoustics and signal processing. His current work is in the area of phased array transducers, and analysis of ultrasonic fields.

Mr. Schafer is a member of Tau Beta Pi, Phi Kappa Phi, and the Acoustical Society of America.

Peter A. Lewin received the M.S. degree in elec­trical engineering in 1968 and the Ph.D. degree in electrical engiiteering and physical acoustics in 1978 in Copenhagen, Denmark.

Before receiving the Ph.D. degree, he was employed by Brue! and Kjaer, Denmark, when he was involved in the development of under­water piezoelectric transducers and associated electronics. From 1978 to 1983, he was asso­ciated with the Danish Institute of Biomedical Engineering and the Technical University of

Denmark, where his research activities primarily focused on propaga­tion of ultrasonic waves in inhomogeneous media. In 1983 he joined the faculty of Drexel University, Philadelphia, PA, in the Department of Electrical Engineering and the Biomedical Engineering and Science Institute. His current research interests are primarily in the field of biomedical ultrasonics including ultrasonic transducer design, image processing, ultrasonic dosimetry, biological effects of ultrasound, and ultrasonic tissue characterization.

Dr. Lewin is a member of the American Institute for Ultrasound in Medicine and the Acoustical Society of America.