Experimental System Prototype of a Portable, Low-Cost, C-Scan Ultrasound Imaging Device

IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 55, NO. 2, FEBRUARY 2008 519

Experimental System Prototype of a Portable,Low-Cost, C-Scan Ultrasound Imaging Device

Michael I. Fuller*, Member, IEEE, Karthik Ranganathan, Member, IEEE, Shiwei Zhou, Member, IEEE,Travis N. Blalock, Member, IEEE, John A. Hossack, Senior Member, IEEE, and William F. Walker, Member, IEEE

Abstract—A system prototype of a future compact, low-costmedical ultrasound device is described and presented with experi-mental results. The prototype system consists of a 32 32 element,fully sampled 2-D transducer array and a printed circuit board(PCB) containing 16 custom “front-end” receive channel inte-grated circuits (ICs) with analog multiplexing and programmablelogic. A PC that included a commercially available data acquisi-tion card is used for data collection and analysis. Beamforming isperformed offline using the direct sampled in-phase/quadrature(DSIQ) algorithm. Pulse-echo images obtained with the prototypeare presented. Results from this prototype support the feasibilityof a low-cost, pocket-sized, C-scan imaging device.

Index Terms—Application-specific integrated circuits, biomed-ical acoustic imaging, C-scan, portable ultrasound.

I. INTRODUCTION

AS DIGITAL beamforming techniques are refined andadvances in semiconductor technology continue to allow

for smaller, more power-efficient, and less expensive integratedcircuits (ICs), the ultrasound community has begun focusingon developing portable medical ultrasound devices [1]–[11].These compact systems allow for the expansion of ultrasoundinto application areas formerly excluded by system size andcost considerations. The newer systems are lightweight (often“hand-held”) units that better facilitate patient point-of-careand offer an order of magnitude cost reduction. Examples ofthese newer, smaller commercial systems include: the iLookseries (Sonosite, Inc., Bothell, WA); OptiGo (Philips MedicalSystems, Andover, MA); Acuson Cypress (Siemens MedicalSolutions USA, Malvern, PA); HS-1500 (Honda Electronics,Co., Ltd., Toyohashi, Aichi, Japan); Terason t3000 (Teratech

Manuscript received February 17, 2007; revised May 16, 2007. This workwas supported by the Carilion Biomedical Institute and NIH NIBIB Grant RO1EB0023489. Asterisk indicates corresponding author.

*M. I. Fuller was with the Department of Biomedical Engineering, Universityof Virginia, Charlottesville, VA 22908 USA. He is now with PocketSonics, Inc.,Charlottesville, VA 22901 USA (e-mail: [email protected]).

K. Ranganathan was with the Department of Biomedical Engineering, Uni-versity of Virginia, Charlottesville, VA 22908 USA. He is now with Pocket-Sonics, Inc., Charlottesville, VA 22901 USA.

S. Zhou was with the Department of Biomedical Engineering, Universityof Virginia, Charlottesville, VA USA. He is now with Philips Research NorthAmerica, Briarcliff Manor, NY 10510 USA.

T. N. Blalock is with the Department of Electrical and Computer Engineering,University of Virginia, Charlottesville, VA 22904 USA.

J. A. Hossack is with the Department of Biomedical Engineering, Universityof Virginia, Charlottesville, VA 22908 USA.

W. F. Walker is with the Department of Biomedical Engineering and the De-partment of Electrical and Computer Engineering, University of Virginia, Char-lottesville, VA 22904 USA.

Digital Object Identifier 10.1109/TBME.2007.903517

Corporation, Burlington, MA); the LOGIQ Book (GE MedicalSystems, Waukesha, WI); and the z.one (ZONARE MedicalSystems, Inc., Mountain View, CA).

While these devices have already found widespread use, avast niche for medical ultrasound remains unfilled among un-conventional or uninitiated users of ultrasound. Such clinicians,medical technicians, battlefield medics, and veterinarians wouldgreatly benefit from using ultrasound to provide adjunct infor-mation during routine medical examinations or when rapid di-agnosis is crucial to patient survival. Adapting the developmentof a medical ultrasound device to these clinical needs requiresjudicious tradeoffs in system complexity to provide an avenueto creating a pocket-sized unit with a simple, intuitive interfaceand order of magnitude cost reduction beyond that offered bycurrent hand-held systems. Furthermore, by employing a 2-Dtransducer array, more intuitive scan formats such as C-modebecome available to users who may have had limited exposure toB-mode imaging. The Sonic Window is a low-cost, pocket-sizedmedical ultrasound system currently under development at theUniversity of Virginia with these concepts and target applica-tions in mind.

Potential applications for the Sonic Window include guidingneedle and catheter insertion [12]–[16]; guiding biopsies [17],[18]; locating foreign bodies [19], [20]; identifying internalbleeding and fluid collection [13], [21]; and supporting routinephysical examination [3], [21]–[24]. The low cost and com-pact size of the device could open applications in veterinarymedicine and animal research. One such application wouldbe tumor localization and growth monitoring, a task that iscurrently performed via palpation or visual inspection. In thisand other applications, the Sonic Window is envisioned as awidely available, easy to use device suitable for users with littleor no prior ultrasound experience. As such, it will not competewith state of the art systems with respect to image quality orflexibility, but must surpass them with respect to ease of use,cost, and portability.

The high level of integration required by the Sonic Windowrequires a significant degree of collaboration and cross-innova-tion between transducer, electronics, and beamforming develop-ment. There exists a need, therefore, to experimentally verify theexpected performance and tradeoffs of these three componentsnot only in isolation, but also as a system, and in a manner versa-tile enough to accommodate modifications in their design. It isprobable that several versions of the transducer will be designedand fabricated as the dicing procedure, backing design, and ma-terial selection are refined. Also, any modifications to the re-ceive electronics circuitry will require new ICs to be fabricated.

0018-9294/$25.00 © 2008 IEEE

520 IEEE TRANSACTIONS ON BIOMEDICAL ENGINEERING, VOL. 55, NO. 2, FEBRUARY 2008

Fig. 1. Long-term concept of the Sonic Window device. A model of the SonicWindow is used to demonstrate its use in guiding needle insertion, one of manypossible applications.

Equally important, the development of an experimental systemcontaining a custom, fully sampled 2-D array and custom re-ceive circuitry provides a unique opportunity to study the per-formance of a variety of beamforming and signal processing al-gorithms while having knowledge and control over design pa-rameters that are inaccessible on commercial clinical scanners.

This paper discusses the design and assembly of an exper-imental ultrasound system consisting of a 32 32-element,fully sampled 2-D transducer array, a PCB containing sixteen64-channel front-end receive ICs with associated control logicand bias circuits, a PCI card digitizer, and a PC that executesbeamforming and signal processing algorithms in software.The 2-D transducer array was fabricated on a separate PCBand mates with the receive electronics PCB by way of surfacemount connectors. This experimental system can capture andstore 1024 channels of RF data for subsequent offline pro-cessing, enabling us to explore the effects of our transducerfabrication, circuit topologies, and beamforming strategieson image quality. A description of the system and obtainedexperimental results are provided next.

II. SYSTEM-LEVEL DESCRIPTION

A description of the proposed ultracompact, low-cost med-ical ultrasound device, the Sonic Window, was presented earlier[25], [26]. The ultimate concept (Fig. 1) is a fully integrated,pocket-sized unit consisting of a 2-D array, receive and protec-tion circuitry implemented on a custom integrated circuit, beam-forming implemented on a digital signal processor (DSP), and ahigh-resolution LCD screen for image display. Crucial savingsin circuit area and complexity are gained through the use of anovel beamforming approach, direct sampled in-phase/quadra-ture (DSIQ) beamforming [27], as well as new integrated circuittopologies [25], [35]. These innovations, combined with the de-velopment of an inexpensive method for fabricating a fully sam-pled 2-D transducer array on a PCB [28], enable an unprece-dented level of integration and dramatic reductions in systemcost.

Fig. 2. System block diagram of the proposed Sonic Window device. The de-sign is partitioned into a single transmit circuit, a fully sampled 2-D transducerarray, a custom IC containing receive circuitry, a commercially available DSP,and a liquid crystal display.

A block diagram of the future Sonic Window system is illus-trated in Fig. 2. The design is partitioned into a single transmitcircuit, a fully-sampled 2-D transducer array, a custom IC con-taining an array of receive channels, a commercially availableDSP, and a liquid crystal display. The common node of the2-D array is connected to the transmit circuit during transmitmode and to analog ground during receive mode. Each receivechannel will consist of an on-chip transmit protection shuntingdevice, a variable gain preamplifier, a bandpass filter, a sample-and-hold (S/H) stage, an analog-to-digital converter (ADC), andstatic memory. Placing the transmit protection devices on-chipeliminates the need for bulky, expensive, and power-consumingoff-chip switching elements. The S/H stage consists of two S/Hunits, whose output samples are one quarter period apart at thecenter frequency of the received pulse. This operation estimatesthe in-phase (I) and quadrature (Q) components of the RF signal,as defined by the DSIQ beamforming algorithm [27]. Further-more, by forming C-mode images rather than B-mode, each S/Hunit need only capture a minimum of one sample per image.The combination of these two properties dramatically simpli-fies the design of the ADC by permitting digitization rates aslow as 10 kHz and produces much less stringent memory anddata bandwidth requirements. Since a standard CMOS processis used, this results in significant reductions in cost, IC area, andpower consumption.

III. DESCRIPTION OF EXPERIMENTAL SYSTEM PROTOTYPE

An experimental ultrasound system was designed and con-structed (Fig. 3) consisting of a 2-D array, custom receiveand protection circuitry, and a PC that included a Gage Com-puscope 12100 PCI card (Gage Applied Technologies, Inc.,Lachine, QC, Canada). The transducer array was fabricated ona PCB substrate [28]. Each transducer element was electricallyconnected to a dedicated pad on a surface mount connector onthe back of the PCB, while all the elements shared a commonconnection on the top. A separate ten-layer, in inPCB was fabricated (Fig. 4) that contains 16 custom receivecircuitry ICs, four ADG707 8:1 differential analog multiplexingICs (Analog Devices, Norwood, MA), two output buffer chan-nels, and an onboard XCR3064XL complex programmablelogic device (CPLD) (Xilinx, Inc., San Jose, CA). The PCBcontaining the 2-D transducer array was designed to matewith the receive circuitry PCB by way of the surface-mount

FULLER et al.: EXPERIMENTAL SYSTEM PROTOTYPE OF A PORTABLE, LOW-COST, C-SCAN ULTRASOUND IMAGING DEVICE 521

Fig. 3. Schematic diagram of the experimental prototype system. The trans-ducer array was fabricated on a PCB substrate. A separate PCB contains 16custom receive circuitry ICs each consisting of 64 front-end receive channels.Each transducer element is electrically connected to a unique receive channel onthe electronics PCB. The outputs of the 1024 receive channels are multiplexedon-chip (64:2) and off-chip (32:2) down to two bandpass filter output channelscoupled to a PC fitted with a two-channel data acquisition card.

connectors (FX11-series, Hirose Electric (USA), Inc., SimiValley, CA). The data acquisition card was capable of capturingtwo channels of 12-bit data simultaneously at a sampling rateof 50 MHz. Thus, 512 separate transmit events were necessaryto capture all 1024 channels of RF data, and memory depthin each channel was limited to 1600 12-bit samples. VolumeRF data acquired with the prototype system was sampled andstored on a PC for subsequent processing and data analysis inMatlab (The MathWorks, Inc., Natick, MA).

A. Fully Sampled, 2-D Transducer Array

Transducer development for the Sonic Window [28] in-volves the dual challenge of constructing a fully sampled (highchannel count) 2-D array that is also inexpensive. The largestdifficulty is the interconnect, where an electrical connectionmust be made between each transducer element and its dedi-cated receive channel. To date, we have used a printed circuitboard substrate, which can be fabricated much less expen-sively than other approaches such as multilayer flex circuits,wire-bonding, and solder connections. The transducer wasfabricated by attaching a 0.53-mm-thick wafer of high per-mittivity, high-electromechanical-coupling-coefficient ceramic(HD3203, CTS Wireless, Albuquerque, NM) to an array of padson the PCB using Chobond silver epoxy (Chomerics, Woburn,MA). The wafer was then diced using a NBC-ZH2040 Discodicing blade (Disco Corp., Tokyo, Japan) to form isolated ele-ments. The kerfs were 0.04 mm wide and 0.125 mm deep intothe PCB substrate. A low-viscosity, unfilled epoxy (RE2039,Loctite, Rocky Hill, CT) was used to fill the kerfs. Finally,the entire transducer array was covered with a gold-platedpolyester electrode that serves as a common node to all theelements [Fig. 5(a)].

The PCB (MetroCircuits, Rochester, NY) consists of ninerouting layers to fan out each element pad to a contact on oneof 16 surface-mount (SM) connectors on the rear of the board[Fig. 5(b)], which electrically mate the transducer array tocorresponding SM connectors on the receive circuitry PCB. Atenth layer serves as a ground plane. Design parameters suchas trace width and spacing (3 mil/76 m), minimum drill size(6 mil/152 m), and pad size (12 mil/305 m) represent thethreshold of current industry capabilities for a ten-layer PCB.

B. Custom-Integrated Circuit Design

A number of researchers have recently implemented front-end receive electronics on custom ICs for applications such asreal-time 3-D imaging [29], intravascular ultrasound [30], [31],intra-oral ultrasound [32], high-frequency annular arrays [33],and portable ultrasound [6], [8], [34]. We designed a custom ICcontaining 64 analog front-end receive channels implemented ina standard TSMC 0.35- m CMOS process available through theMOSIS Integrated Circuit Fabrication Service (Marina del Rey,CA). Sixteen of these 64-channel ICs are used to form all 1024receive channels in our prototype system. Each channel consistsof an on-chip transmit protection shunting device, a variablegain preamplifier, and a transconductance buffer. The receivechannel is fully differential to reduce distortion and suppressesthe effects of power supply and substrate noise.

The transmit protection scheme, described in [35], excludesthe expensive and area-consuming off-chip components usedin other systems to prevent the high-voltage transmit pulsefrom damaging the receive electronics. Instead, a suitablysized NMOS transistor implemented on-chip is connectedbetween the preamplifier input and a low-impedance powersupply serving as analog ground. During the transmit event,this NMOS transistor is turned on, shunting the large currenttransient from the high-voltage transmit pulse to analog ground.Only a fraction (on the order of 100 mV) of the transmit voltage(as high as 100 V) appears at the input to the preamplifier. Theshunt device is turned off during receive to permit amplificationof the received echo signal.

The low-noise preamplifier design (Fig. 6) consists of twoidentical differential stages with variable gain. The gain can beadjusted between 30 and 85 dB by adjusting the bias voltageof triode-region device M3, which serves as a source degen-eration resistance (Fig. 7). A novel low-frequency suppressionscheme is incorporated into the preamplifier design to serve thedual purpose of reducing 1/f noise and rejecting dc offset. Theactive load of the amplifier is designed to have high mid-bandgain, but small low-frequency gain, as shown in Fig. 7. The gainprofile can be tuned by adjusting the bias voltage LowFreqAdj.The preamplifier equivalent input noise of 5 nV/ Hz (withinthe band of interest) was found in simulation by performing anoise analysis in the Cadence Virtuoso Spectre Circuit Simu-lator (Cadence Design Systems, Inc., San Jose, CA).

Each preamplifier is followed by a differential transconduc-tance buffer performing a voltage-to-current conversion, pro-viding the means to implement a current-mode analog multi-plexing scheme in which multiple channels can share the sameoutput node with minimal impact on signal bandwidth. The 64channels are grouped into two 32-channel banks each having a


Fig. 4. Photograph of the 11 in x 11.5 in receive electronics PCB (top view). The PCB contains sixteen 64-channel custom receive circuitry ICs, four differentialanalog multiplexing ICs, two output buffer channels, and an onboard CPLD. The 2-D transducer array was designed to mate with the rear side of the receivecircuitry PCB by way of surface-mount connectors.

differential output (Fig. 8). One channel per bank is permittedto drive its signal onto the common output nodes at any giventime. The output buffers of the inactive channels are disabledby turning off their respective output MOSFET devices, effec-tively forcing them into a high-impedance state. A 5-bit decoderselects the active channel in each bank.

The active die area of the front-end receive circuit IC is1.9 mm 0.9 mm, and includes the 64 analog receive chan-nels and associated multiplexing logic (Fig. 9). The die werepackaged (Promex Industries, Santa Clara, CA) into a KyoceraPGA121M (Kyocera America, Inc., San Diego, CA) 121-pinceramic package.

IV. METHODS

The experimental prototype system was assembled and testedfor basic functionality prior to attempting a transmit event orforming pulse-echo images. Test points on the back of the trans-ducer array were probed with the transducer PCB connected tothe receive electronics PCB. A 100-mV sinusoidal test signalwas applied to each element and RF data were acquired from

all channels. This experiment provided verification of the map-ping of transducer elements to channels and identified open orshorted connections between the transducer elements and theinputs to the receive circuitry ICs.

Pulse-echo volume datasets were acquired using the proto-type system. The transducer was driven at its center frequencyof 3.3 MHz with an eight-cycle Gaussian-enveloped sinusoidhaving a full-width at half-maximum (FWHM) bandwidth ofapproximately 30% and a peak-to-peak amplitude of 30 V. Asdescribed above, the data acquisition hardware was capable ofacquiring and storing 1024 channels of data sampled at 50 MHzwith a memory depth of 1600 12-bit samples. C-mode images ofeach target were formed on the PC in Matlab using three beam-forming techniques: beamforming using only time delays, con-ventional baseband demodulated I/Q beamforming using onlyphase delays, and DSIQ beamforming that also uses only phasedelays. Beamforming was implemented to be consistent withthe methods followed in [27], with the exception of the assumedtransducer fractional bandwidth (30% versus 55%), center fre-quency (3.3 MHz versus 5.5 MHz), the sampling rate (50 MHzversus 39.27 MHz), and that fact that our experimental system


Fig. 5. Fully sampled 2-D transducer array PCB. The transducer array is lo-cated on the top of the PCB (a) where the gold foil common node is visible.The surface-mount connectors on the bottom of the PCB (b) interface with thereceive electronics PCB and connect each transducer element to its respectivereceive channel.

forms C-mode images (in keeping with the Sonic Window de-vice concept) as opposed to the B-mode images formed in [27].The acquired data was filtered in Matlab to 30% -dB frac-tional bandwidth at 3.3 MHz with a 51-order FIR filter. Hannwindow apodization was used [36]. Scaling was applied whereapplicable to compensate for energy differences in pixels beam-formed with receive apertures that intersected edges of the 2-Darray.

Time-delay beamforming was implemented using unquan-tized time delays (IEEE double precision floating point repre-sentation). A cubic spline-based continuous representation ofthe sampled data was used to evaluate the received signals ineach channel at the requisite time points [27], [37]. RF dataproduced by summing across channels were envelope detectedusing the Hilbert transform.

Fig. 6. Schematic of a single stage of the differential preamplifier. The gaincan be adjusted from 30 to 85 dB by adjusting the bias voltage RecvGain. Theactive load of the amplifier is designed to have high mid-band gain, but smalllow-frequency gain. The gain profile can be tuned by adjusting the bias voltageLowFreqAdj.

Fig. 7. Experimental measurements of gain (at 5 MHz) as a function of Recv-Gain control voltage and frequency response of preamplifier at maximum gainsetting (RecvGain = 2 V).


Fig. 8. Block diagram of the 64-channel receive circuitry front-end IC. Eachchannel is fully differential and consists of an NFET protection device, a pream-plifier, and a transconductance output buffer. The channels are arranged in two32-channel banks. All channels within a bank share a differential output node.A 5-bit decoder selects one channel per bank at a time by enabling the channel’soutput buffer.

Fig. 9. Microphotograph of the 64-channel receive circuitry IC [35].

Conventional baseband demodulated data were obtainedby multiplying the received data by to formthe in-phase (I) component and by to form thequadrature (Q) component, where was the transducer centerfrequency of 3.3 MHz, was the sample number, and wasthe sampling interval. This data were then low-pass filteredwith a fifth-order Butterworth filter and zero-phase distortionwas accomplished by filtering once in the forward directionand then filtering a second time after reversing the output. Thisproduced a -dB cutoff at 3.3 MHz. The I and Q componentswere combined into an analytic representation of the receivedecho signal that was then apodized and focused via phaserotation through complex multiplication operations, the resultsof which were summed across channels to yield the intensity ata given point on the image plane.

As described in [27], DSIQ samples will be formed in ourfinal system by splitting the received signal between two parallelsample-and-hold (S/H) channels. As described in Section II, theSonic Window device will ultimately include S/H stages andADCs on the same IC as the front-end receive circuitry com-ponents. In each receive channel, one S/H will acquire the Icomponent and then the other S/H will acquire the Q compo-nent by sampling one quarter of a period later at the centerfrequency of the signal. In the experimental system prototype,only the front-end circuitry was implemented on-chip, and thesampling was performed by the Gage Compuscope 12100 PCIcard at a rate of 50 MHz at uniform sampling intervals. TheI component sample was taken directly from this acquired RFdata—the Q component sample was synthesized by interpo-lating the acquired RF data at a time lag of a quarter periodat the assumed center frequency using cubic spline interpola-tion [27], [37]. The apodization, phase rotation and summationacross channels were implemented as in the conventional base-band demodulation case described above.

The three beamforming methods described above were usedto form C-mode images from pulse-echo data acquired off twotargets for the purpose of comparing the performance of theDSIQ method to the time-delay and conventional I/Q methods,as well as evaluate the overall performance of the experimentalsystem. The first target was a 200- m nylon wire in a watertank placed 1.5 cm below and parallel to the face of the trans-ducer. The second target was a custom-made “edge phantom,”which consisted of 10% acrylamide gel ( m/s) havingone speckle-generating region and one nonspeckle-generatingregion. The speckle-generating region was constructed by incor-porating Sephadex (Amersham, Piscataway, NJ) into the acry-lamide. The procedure followed in constructing this phantomwas based on that described in [38], although higher acrylamideconcentrations were used. The geometry of this edge phantomis such that an ideal C-mode image acquisition should producean image with speckle in one half and an absence of speckle(anechoic region) in the other half.

V. EXPERIMENTAL RESULTS

Fig. 10 is a binary mapping of “dead” channels discoveredfrom the probing procedure. It was found that 69 out of 1024electrical channels (6.74%) contained an open circuit betweenthe transducer element and its corresponding receive channelinput. The cause of the majority of these open connectionswas determined to be due to errors in the layout design ofthe 2-D transducer array PCB. Poor contact was also notedbetween the surface-mount connectors on the transducer PCBand the receive electronics PCB. Since contact between thesesurface-mount connectors relies on a friction fit, minor bendingand warping of the transducer PCB can result in localizedmisalignment of connector contacts. This overall net channelyield should be distinguished from the transducer elementyield, which was measured to be 99%.

The top panel in Fig. 11 illustrates the simulated 2-D pointspread function (PSF) in the C-mode plane for the ideal case inwhich 100% of the receive channels are connected to their re-spective transducer elements. The bottom panel in Fig. 11 illus-trates the simulated 2-D PSF in the C-mode plane for a 6.74%


Fig. 10. Diagram mapping the location of dead channels (defined as an opencircuit between the receive channel and transducer element, indicated in black)with respect to their corresponding physical location on the 2-D transducerarray.

channel loss having the same distribution (channel-to-elementmapping) observed in the experimental prototype system (seeFig. 10). These two PSFs were computed in Matlab using anarrowband Rayleigh–Sommerfeld formulation [39] for a planeparallel to the transducer at a depth of 2 cm. Hann windowapodization [37] was used. The images in Fig. 11 were normal-ized, logarithmically compressed, and mapped so as to presentimage intensity over the range to 0 dB.

The C-mode images of the wire target acquired with theprototype system (Fig. 12) were formed using conventionalunquantized time-delay beamforming (TD), conventionalbaseband demodulated I/Q beamforming (IQ), and DSIQbeamforming (DSIQ) methods for f#s of 0.5, 1, and 2. Hannwindow apodization was used. The C-mode images werenormalized, logarithmically compressed, and mapped so asto present image intensity over the range to 0 dB. TheFWHM of the wire in the acquired images at f/1 is 1.2 mm.One-dimensional integrated cross-sections of the wire targetimages in azimuth were formed by performing a 2-D splineinterpolation (8x) on the original absolute image, summing inelevation, and then plotting the normalized, logarithmicallycompressed result.

C-mode images of the edge phantom target acquired with theprototype system (Fig. 13) were formed using pure time-delaybeamforming (TD), conventional baseband demodulated I/Qbeamforming (IQ), and DSIQ beamforming (DSIQ) methodsfor f#s of 0.5, 1, and 2. Hann window apodization was used.The contrast between the region with speckle and region withno speckle was 10 dB. The C-mode images were normalized,logarithmically compressed and mapped so as to present imageintensity over the range to 0 dB. One-dimensional inte-grated cross-sections of the edge phantom images in elevationwere formed by performing a 2-D spline interpolation (8x) on

Fig. 11. Simulated 2-D PSF (C-mode plane at a depth of 2 cm) for the idealcase in which 100% of the receive channels are connected to their respectivetransducer elements, and for a 6.74% channel loss following the distribution(channel-to-element mapping) observed in the experimental prototype system.Both images were normalized, logarithmically compressed and mapped so as topresent image intensity over the range �40 to 0 dB.

the original absolute image, summing in azimuth, and thenplotting the normalized, logarithmically compressed result.

VI. DISCUSSION

Overall, the experimental results shown are very promisingin that they demonstrate successful pulse-echo image formationwith a low-cost, compact, experimental ultrasound system in itsproof-of-concept phase of development. There do exist, how-ever, multiple avenues for improving image quality.

The largest contributor to poor image quality was the pres-ence of large grating lobes (observable in Fig. 11 approximately1.5 cm from the focus) caused by the transducer element pitch.The pitch was 635 m, while the wavelength at 3.3 MHz and1540 m/s speed of sound is 467 m. Element pitch was limitedby the PCB manufacturing capabilities, specifically the tracewidth/spacing and minimum pad size for vias. The minimumpad size was necessarily increased to account for the worseningdrill tolerance as the aspect ratio (ratio of PCB thickness to


Fig. 12. C-mode images acquired with the prototype system of a 200-�m nylon wire in a water tank placed 1.5 cm below and parallel to the face of the transducer.The images were formed using pure time-delay beamforming (TD), conventional baseband demodulated I/Q beamforming (IQ), and DSIQ beamforming (DSIQ)methods for f#s of 0.5, 1, and 2. Hann window apodization was used. The images were normalized, logarithmically compressed, and mapped so as to present imageintensity over the range �30 to 0 dB. One-dimensional integrated cross-sections of the wire target images in azimuth were formed by performing a 2-D splineinterpolation (8x) on the original absolute image, summing in elevation, and then plotting the normalized, logarithmically compressed result.

drill hole size) increased. The aspect ratio increased with thenumber of routing layers, which in turn depended on the tracewidth/spacing.

Another significant contributor to image quality limitations isthe manner in which signals were routed from individual trans-ducer elements to their respective electronic receive channels.As described in Section V, routing errors in the transducer PCBalong with intermittent poor contact between the surface mountconnectors resulted in a channel loss of 6.74%. As illustrated inFig. 11, these lost channels have the undesirable effect of signifi-cantly degrading the PSF by raising side lobe levels to as high as

dB and distorting the mainlobe. Note also that the channelloss results in a shift variant system—pixel-to-pixel gain variesas the receive aperture is translated across this nonuniform pat-tern of “dead” transducer elements. These effects significantlycontribute to the presence of clutter and other artifacts, causinganomalous spots and kinks or gaps in the wire target image andpoor contrast in the edge phantom image. Furthermore, the inter-

connect scheme also suffers from parasitic inductances, capaci-tances, and resistances associated with long PCB traces, the sur-face-mount connectors, and the custom IC pin grid array (PGA)chip packages that all contribute to signal-to-noise (SNR) degra-dation, crosstalk, and impedance variation among channels.

Despite successful image formation using this preamplifier(see Fig. 6), a significant problem was discovered involvingthe low-frequency suppression scheme. When a large transientvoltage appeared at the input of the preamplifier (similar to thataccompanying a transmit event), the drain-source voltage of thetriode-region device M8 (M9) was large enough that a chargewas drained from the MOScap, M12 (M13), in the active load.This changed the bias point enough to significantly lower theoverall mid-band gain of the preamplifier. A temporary solutionwas implemented in which the LowFreqAdj bias was dynami-cally tuned such that after each transmit event some of the chargewas allowed to return onto M12 (M13), though the gain anddynamic range of the overall preamplifier was still degraded.


Fig. 13. C-mode images acquired with the prototype system of a custom-made “edge phantom,” which consisted of 10% acrylamide gel (c = 1545 m/s) having onespeckle-generating region and one nonspeckle-generating region. The geometry of this edge phantom is such that an ideal C-mode image acquisition should producean image with speckle in the bottom half and an absence of speckle (anechoic region) in the top half. The images were formed using pure time-delay beamforming(TD), conventional baseband demodulated I/Q beamforming (IQ), and DSIQ beamforming (DSIQ) methods for f#s of 0.5, 1, and 2. Hann window apodizationwas used. The images were normalized, logarithmically compressed, and mapped so as to present image intensity over the range �20 to 0 dB. One-dimensionalintegrated cross sections of the edge phantom images in elevation were formed by performing a 2-D spline interpolation (8x) on the original absolute image,summing in azimuth, and then plotting the normalized, logarithmically compressed result.

Another factor influencing image quality is the use of theDSIQ beamforming algorithm, which makes concessions inimage quality to provide the dramatic hardware savings ex-ploited by the Sonic Window device. However, as illustratedin Figs. 12 and 13, the C-mode images formed using DSIQbeamforming exhibit very little difference from the conven-tional time-delay and I/Q methods. While it can be arguedthat the transducer, interconnect, and front-end shortcomingsdescribed above dominate the subtler differences in imagequality that exist between the three beamforming approaches,the integrated 1-D cross sections in Figs. 12 and 13 offerfurther insight. Here, the differences between beamformingapproaches are more evident—particularly at lower f#s—butthe overall performance is still similar. DSIQ and conventionalI/Q beamforming appear to behave almost identically, andboth deviate from TD beamforming at lower f#s because both

approaches rely on phase rotation to achieve delays and are thussusceptible to focusing errors toward the edge of the aperture.The integrated 1-D cross-sections at higher f#s demonstratecloser agreement among all the beamforming approaches.These trends are similar to the findings in [27], where the DSIQalgorithm was shown to compare favorably with conventionalbeamforming techniques in a B-mode commercial ultrasoundscanner not suffering from the grating lobe and channel-lossissues present in our experimental system. DSIQ beamformingwas also shown to be surprisingly robust to error in the assumedcenter frequency, which would cause the S/H clocks to beoffset at phase differences other than 90 . According to Fig. 6in [27], an 18% error in the assumed center frequency (70phase difference between I and Q samples) led to a loss of only1 dB in contrast performance. Additionally, while narrowbandsignals are the ideal for DSIQ beamforming, reasonable image


quality was obtained with fractional bandwidths as high as55% [27].

Ultimately, the target clinical application for the SonicWindow device dictates the losses in image quality that can betolerated in exchange for the low cost, compact size, and simpleoperation that it demands. A consequence of the transmitprotection scheme is the lack of focusing on transmit and its as-sociated losses in SNR and penetration. However, as discussedin [35], the Sonic Window is intended for imaging shallowtargets, which significantly relaxes the penetration requirement,and all image data is captured in parallel after each transmitevent, so a plane wave transmit is appropriate. Furthermore,the protection scheme significantly simplifies the transmitterhardware [27], facilitates a flip-chip interconnect scheme infuture versions of the Sonic Window (avoiding SNR lossesthrough cabling) [35], and is critical in enabling the use of a2-D array, which provides array gain benefits and compensatesfor slightly worse clutter in azimuth with much better perfor-mance in elevation [40]. Similarly, a consequence of DSIQbeamforming is its preference for a smaller signal bandwidth,which corresponds to an increase in C-plane thickness. Ourexperimental system used a 3.3-MHz center frequency and 30%fractional bandwidth generating a slice thickness of approxi-mately 770 m. However, this is more than adequate to satisfyour initial goal of visualizing a 5-mm-diameter blood vessel inthe arm for needle or IV line insertion [40] and slice thicknesswill decrease as the system center frequency increases in futureprototypes.

The experimental system described in this paper servedwell as confirmation of our general design approach, noveltechnologies, and performance tradeoffs. Several modificationswill be necessary to realize the long-term concept of a clinicallyviable Sonic Window device. The grating lobe and channelloss issues can be addressed through an improved interconnectapproach providing smaller transducer element pitch and in-creased channel count. By designing receive electronics ICswith a channel pitch that corresponds to the transducer elementpitch, a flip-chip interconnect scheme is possible, providing astraightforward connection between transducer elements andreceive channels with little or no routing [8], [30], [31], [41],[42]. If DSIQ beamforming is used, digitization and memorycan be brought on-chip along with the front-end receive elec-tronics, enabling the simultaneous capture of all the complexdata needed to form a C-scan image—a simple depth controlwould vary the timing of the global S/H clock signals. Ourexperimental system took as long as two minutes to form oneC-scan image, which included time for 512 transmits, memorytransfer of the volume data from the acquisition card, andprocessing on the PC. However, with parallel on-chip datacapture, DSIQ beamforming could be performed in real-timeby a low cost, commercially available DSP [27], [42]. Thesignificantly low pulse repetition frequency (as low as theframe rate) requirements of such an approach can be exploitedby switching off the transmit and receive electronics in betweentransmit events to save power and prevent overheating, evenallowing for multiple transmit events for signal averaging toimprove SNR [35] and compounding techniques for specklereduction [27].

VII. CONCLUSION

A high-channel-count experimental ultrasound system wasconstructed and experimentally shown to successfully formpulse-echo images of a wire target and edge phantom. Imagesformed using the DSIQ beamforming algorithm compared fa-vorably with that of conventional time-delay beamforming andbaseband demodulated I/Q beamforming. Image quality wasimpacted predominantly by grating lobes caused by suboptimaltransducer element spatial sampling and routing errors in thetransducer PCB coupled with poor surface mount connectorcontact. Future efforts will focus on reducing element pitchand utilizing alternative interconnect approaches to providesignificant improvements in image quality in the future SonicWindow system.

ACKNOWLEDGMENT

The authors would like to thank E. Girard for her contributionto the transducer design, E. Brush for his assistance in probingthe transducer interconnect, and M. Oberhardt for constructingour phantoms.

REFERENCES

[1] L. D. Greenbaum, “It is time for the sonoscope,” J. Ultrasound Med.,vol. 22, no. 4, pp. 321–322, 2003.

[2] R. A. Filly, “Is it time for the sonoscope? If so, then let’s do it right!,”J. Ultrasound Med., vol. 22, no. 4, pp. 323–325, 2003.

[3] J. R. T. C. Roelandt, “Ultrasound stethoscopy: A renaissance of thephysical examination?,” Heart, vol. 89, pp. 971–974, 2003.

[4] S. L. P. Langlois, “Portable ultrasound on deployment,” ADF Health,vol. 4, pp. 77–80, Sep. 2003.

[5] D. D. Price, S. R. Wilson, and T. G. Murphy, “Trauma ultrasound fea-sibility during helicopter transport,” J. Air Med., vol. 19, no. 4, pp.144–146, Oct.–Dec. 2000.

[6] J. Hwang, J. Quistgaard, J. Souquet, and L. A. Crum, “Portable ultra-sound device for battlefield trauma,” in Proc. IEEE Ultrason. Symp.,1998, vol. 2, pp. 1663–1667.

[7] Y. Saijo, S.-I. Nitta, K. Kobayashi, H. Arai, and Y. Nemoto, “Devel-opment of an ultra-portable echo device connected to USB port,” Ul-trasonics, vol. 42, pp. 699–703, Apr. 2004.

[8] T. White, K. Eriksen, and A. Nicoli, “Three-dimensional ultrasonicimaging with a fully populated 128� 128 array,” in Proc. 19th Annu.Int. Conf. IEEE Eng. Med. Biol. Soc., 1997, vol. 2, pp. 744–746.

[9] J. J. Kim and T. K. Song, “An ultrasound beamforming method using1.5-bit ADCs for portable ultrasound scanners,” in Proc. IEEE Utrason.Symp., 23–27 Aug. 2004, vol. 3, pp. 1722–1724.

[10] M. Karaman and M. Donnell, “Synthetic aperture imaging for smallscale systems,” IEEE Trans. Ultrason., Ferroelectr., Freq. Control.,vol. 42, no. 3, pp. 429–442, May 1995.

[11] R. Fisher, K. Thomenius, R. Wodnicki, R. Thomas, S. Cogan, C.Hazard, W. Lee, D. Mills, B. Khuri-Yakub, and A. Ergun, “Reconfig-urable arrays for portable ultrasound,” in Proc. IEEE Ultrason. Symp.,18–21 Sep. 2005, vol. 1, pp. 495–499.

[12] J. M. Rothschild, “Ultrasound guidance of central vein catheterization,”ch. 21, 2001 [Online]. Available: http://www.ahrq.gov/clinic/ptsafety/,Evidence Report/Technology Assessment, No. 43. Ch. 21. MakingHealthcare Safer. A Critical Analysis of Patient Safety Practices.Agency for Healthcare Research and Quality Publication, No. 01-E058.

[13] Y. Beaulieu and P. E. Marik, “Bedside ultrasonography in the ICU:Part 1,” Chest, vol. 128, pp. 881–895, Aug. 2005.

[14] P. D. Levin, O. Sheinin, and Y. Gozal, “Use of ultrasound guidance inthe insertion of radial artery catheters,” Critical Care Med., vol. 31, pt.2, pp. 481–484, Feb. 2003.

[15] A. G. Randolph, D. J. Cook, C. A. Gonzalez, and C. G. Pribble, “Ultra-sound guidance for placement of central venous catheters: A meta-anal-ysis of the literature,” Crit. Care Med., vol. 24, pp. 2053–2058, 1996.

[16] W. R. Fry, G. C. Clagett, and P. T. O’Rourke, “Ultrasound-guided cen-tral venous access,” Archives Surgery, vol. 134, no. 7, pp. 738–741,1999.

[17] N. Velez, D. E. Earnest, and E. D. Staren, “Diagnostic and interven-tional ultrasound for breast disease,” Amer. J. Surgery, vol. 180, no. 4,pp. 284–287, 2000.


[18] A. Liebeskind, A. G. Sikora, A. Komisar, D. Slavit, and K. Fried,“Rates of malignancy in incidentally discovered thyroid nodulesevaluated with sonography and fine-needle aspiration,” J. UltrasoundMed, vol. 24, pp. 629–634, 2005.

[19] H. T. Harcke, A. D. Levy, and G. J. Lonergan, “The sonographic ap-pearance and detectability of nonopaque and semiopaque materials ofmilitary origin,” Military Med., vol. 167, no. 6, pp. 459–463, Jun. 2002.

[20] R. Hill, R. Conron, P. Greissinger, and M. Heller, “Ultrasound for thedetection of foreign bodies in human tissue,” Ann. Emergency Med.,vol. 29, no. 3, pp. 353–356, Mar. 1997.

[21] Y. Beaulieu and P. E. Marik, “Bedside ultrasonography in the ICU:Part 2,” Chest, vol. 128, pp. 1766–1781, 2005.

[22] J. Ulrich and C. Voit, “Ultrasound in dermatology. Part II. Ultrasoundof regional lymph node basins and subcutaneous tumours,” Eur. J. fDermatol., vol. 11, no. 1, pp. 73–79, 2001.

[23] M. L. Ghirardelli, V. Jemos, and P. G. Gobbi, “Diagnostic approach tolymph node enlargement,” Haematologica, vol. 84, pp. 242–247, 1999.

[24] M. T. Reeder, B. H. Dick, J. K. Atkins, A. B. Pribis, and J. M. Martinez,“Stress fractures. Current concepts of diagnosis and treatment,” SportsMed., vol. 22, no. 3, pp. 198–212, Sep. 1996.

[25] M. I. Fuller, T. N. Blalock, J. A. Hossack, and W. F. Walker, “Aportable, low-cost, highly integrated, 3D medical ultrasound system,”in Proc. IEEE Ultrason. Symp., Oct. 2003, vol. 1, pp. 38–41.

[26] M. I. Fuller, K. Ranganathan, S. Zhou, T. N. Blalock, J. A. Hossack,and W. F. Walker, “Portable, low-cost medical ultrasound device pro-totype,” in Proc. IEEE Ultrason. Symp., 23–27 Aug. 2004, vol. 1, pp.106–109.

[27] K. Ranganathan, M. K. Santy, T. N. Blalock, J. A. Hossack, and W. F.Walker, “Direct sampled I/Q beamforming for compact and very low-cost ultrasound imaging,” IEEE Trans. Ultrason., Ferroelectr., Freq.Control., vol. 51, no. 9, pp. 1082–1094, Sep. 2004.

[28] E. Girard, S. Zhou, W. Walker, T. Blalock, and J. Hossack, “High ele-ment count two dimensional transducer array,” in Proc. IEEE Ultrason.Symp., 5–8 Oct. 2003, vol. 1, pp. 964–967.

[29] J. Morizio, S. Guhados, J. Castellucci, and O. von Ramm,“64-channel ultrasound transducer amplifier,” in Proc. SouthwestSymp. Mixed-Signal Design, Feb. 2003, pp. 228–232.

[30] I. O. Wygant, X. Zhuang, D. T. Yeh, S. Vaithilingam, A. Nikoozadeh,O. Oralkan, A. S. Ergun, M. Karaman, and B. T. Khuri-Yakub, “Anendoscopic imaging system based on a two-dimensional CMUT array:Real-time imaging results,” in Proc. IEEE Ultrason. Symp., Sep. 2005,vol. 2, pp. 792–795.

[31] I. Çiçek, A. Bozkurt, and M. Karaman, “Design of a front-end in-tegrated circuit for 3D acoustic imaging using 2-D CMUT arrays,”IEEE Trans. Ultrason., Ferreolectr., Freq. Control, vol. 52, no. 12, pp.2235–2241, Dec. 2005.

[32] L. L. Lay, S. J. Carey, and J. V. Hatfield, “Pre-amplifier arrays for intra-oral ultrasound probe receiving electronics,” in Proc. IEEE Ultrason.Symp., Aug. 2004, vol. 3, pp. 1753–1756.

[33] J. R. Talman and S. L. Garverick, “Integrated-circuit implementation ofa matched-cell dynamic focusing architecture for a 5-channel, 50-MHzplanar annular array,” in Proc. IEEE Ultrason. Symp., Oct. 2001, vol.2, pp. 1109–1112.

[34] M. Sawan, R. Chebli, and A. Kassem, “Integrated front-end receiverfor a portable ultrasonic system,” Analog Integrated Circuits SignalProcess., vol. 36, pp. 57–67, 2003.

[35] M. I. Fuller, T. N. Blalock, J. A. Hossack, and W. F. Walker, “Noveltransmit protection scheme for ultrasound systems,” IEEE Trans. Ul-trason., Ferreolectr., Freq. Control, vol. 54, no. 1, pp. 79–86, Jan. 2007.

[36] F. J. Harris, “On the use of windows for harmonic analysis with thediscrete Fourier Transform,” Proc. IEEE, vol. 66, no. 1, pp. 51–83, Jan.1978.

[37] F. Viola and W. F. Walker, “A spline-based algorithm for continuoustime-delay estimation using sampled data,” IEEE Trans. Ultrason.,Ferreolectr., Freq. Control, vol. 52, no. 1, pp. 80–93, Jan. 2005.

[38] L. A. Negron, F. Viola, E. P. Black, C. A. Toth, and W. F. Walker,“Development and characterization of a vitreous mimicking materialfor radiation force imaging,” IEEE Trans. Ultrason., Ferreolectr., Freq.Control, vol. 49, no. 11, pp. 1543–1551, Nov. 2002.

[39] J. W. Goodman, Introduction to Fourier Optics, 2nd ed. New York:McGraw-Hill, 1988, pp. 46–50.

[40] K. Ranganathan and W. F. Walker, “Cystic resolution: A performancemetric for ultrasound imaging systems,” IEEE Trans. Ultrason., Ferro-electr., Freq. Control, vol. 54, no. 4, pp. 782–792, Apr. 2007.

[41] M. I. Fuller, E. V. Brush, M. D. C. Eames, T. N. Blalock, J. A. Hos-sack, and W. F. Walker, “The sonic window: Second generation pro-totype of low-cost, fully integrated, pocket-sized medical ultrasounddevice,” in Proc. 2005 IEEE Ultrason. Symp., 18–21 Sep. 2005, vol. 1,pp. 273–276.

[42] W. F. Walker, M. I. Fuller, E. V. Brush, M. D. C. Eames, K. Owen, K.Ranganathan, T. N. Blalock, and J. A. Hossack, “The sonic window:Second generation results,” Proc. SPIE, vol. 6147, pp. 97–103, Mar.2006.

Michael I. Fuller (S’01–M’06) received the B.S. andM.S. degrees in electrical engineering from the Uni-versity of Virginia, Charlottesville, in 2001 and thePh.D. degree in biomedical engineering, Universityof Virginia in 2007.

After completing his doctoral work, he joinedPocketSonics, Inc.—a startup devoted to low-costand handheld ultrasound systems—where he iscurrently a Research Scientist and involved insystem and integrated circuit development. He isalso a Visiting Research Scientist at the University

of Virginia. His research interests include analog and mixed-signal integratedcircuit design, ultrasound imaging, and medical devices.

Karthik Ranganathan (S’00–M’05) received theB.E. degree in biomedical engineering from theUniversity of Mumbai, Mumbai, India, in 1999 andthe Ph.D. degree in biomedical engineering from theUniversity of Virginia, Charlottesville, in 2005. Hisresearch interests during his doctoral work includedultrasound beamforming, signal processing, andangular scatter measurement techniques, and hisdissertation explored the two extremes—optimalbeamforming for the best attainable image qualityin state of the art systems and beamforming for very

low-cost systems.After his doctoral work, he joined PocketSonics, Inc.—a startup devoted to

low-cost and handheld ultrasound systems—where he is currently a ResearchScientist and involved in system and beamformer development. He is also aVisiting Research Scientist at the University of Virginia.

Shiwei Zhou (S’03–M’06) was born in Beijing,China, in 1974. He received the B.S. and M.S.degrees in optical-electrical engineering from theBeijing Institute of Technology, Beijing, in 1996 and1999, respectively, and the Ph.D. degree in biomed-ical engineering from the University of Virginia,Charlottesville, in 2005.

He joined the Biomedical Engineering De-partment, University of Virginia, in 2000. He iscurrently a Research Scientist at Philips ResearchNorth America, Briarcliff Manor, NY. His research

interests are finite-element analysis (FEA) modeling for various ultrasoundtransducers and applications, new transducer techniques and optimization,high-intensity focused ultrasound, and digital signal processing techniques inultrasound.

Travis N. Blalock (M’91) received the B.S. and M.S.degrees from the University of Tennessee, Knoxville,in 1985 and 1988, respectively, and the Ph.D. degreefrom Auburn University, Auburn, AL, in 1991 underthe direction of R. Jaeger.

In 1991, he joined Agilent Laboratories, PaloAlto, CA, (formerly Hewlett-Packard Labs) wherehe was involved in several mixed-signal CMOSdesign efforts, including disk read channel filters, amassively parallel analog cross-correlation processorwith integrated image capture, and high-resolution

color liquid crystal on silicon microdisplays. He joined the faculty of theDepartment of Electrical Engineering, University of Virginia, Charlottesville,in 1998. He is currently leading a mixed-signal CMOS research effort at theUniversity of Virginia focusing on integrated signal processing and imaging inmedical ultrasound and biotelemetry.

Dr. Blalock is a member of Tau Beta Pi and Eta Kappa Nu.


John A. Hossack (S’90–M’92–SM’02) was bornin Glasgow, U.K., in 1964. Hereceived the B.Eng.Hons(I) degree in electrical and electronic engi-neering from Strathclyde University, Glasgow, in1986 and the Ph.D. degree in electrical and electronicengineering in 1990. From 1990 to 1992, he was aPostdoctoral Researcher in the E. L. Ginzton Lab-oratory, Stanford University, Stanford, CA, workingunder B. A. Auld’s guidance. His research was onmodeling of 0:3 and 1:3 piezoelectric compositetransducers. In 1992, he joined Acuson, Mountain

View, CA, initially working on transducer design. During his time at Acuson,his interests diversified into beamforming and 3-D imaging. Dr. Hossack wasmade a Fellow of Acuson for “excellence in technical contribution” in 1999.In 2000, he joined the Biomedical Engineering Department, University ofVirginia, Charlottesville. His current research interests relate to transducerdesign, 3-D ultrasound imaging, and contrast microbubble usage in ultrasoundfor imaging and therapy.

William F. Walker (S’95–M’96) received the B.S.E.and Ph.D. degrees from Duke University, Durham,NC, in 1990 and 1995, respectively. His dissertationexplored fundamental limits on the accuracy of adap-tive ultrasound imaging.

After completing his doctoral work, he stayed onat Duke as an Assistant Research Professor in theDepartment of Biomedical Engineering. At the sametime, he served as a Senior Scientist and President ofNovaSon Corporation, Durham. In 1997 he joinedthe faculty of the Department of Biomedical Engi-

neering, University of Virginia, Charlottesville, being promoted to AssociateProfessor in 2003. He is an active founder in two ultrasound-based startupcompanies in Charlottesville: PocketSonics, Inc. and HemoSonics, LLC. Hisresearch interests include aperture domain processing, beamforming, angularscatter imaging, tissue elasticity imaging, low-cost system architectures, andtime delay and motion estimation.

Documents

Experimental System Prototype of a Portable, Low-Cost, C-Scan Ultrasound Imaging Device