Hansen, Jens Munk

Hemmsen, Martin Christian

Jensen, Jørgen Arendt

Synthetic Aperture Sequential Beamforming (SASB) is applied to medical ultrasound imaging using a multi element convex array transducer. The main motivation for SASB is to apply synthetic aperture techniques without the need for storing RF-data for a number of elements and hereby devise a system with a reduced system complexity. Using a 192 element, 3.5 MHz, λ-pitch transducer, it is demonstrated using tissue-phantom and wire-phantom measurements, how the speckle size and the detail resolution is improved compared to conventional imaging

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Downloaded from orbit.dtu.dk on: Dec 20, 2017Synthetic Aperture Sequential Beamformation applied to medical imaging.Hemmsen, Martin Christian; Hansen, Jens Munk; Jensen, Jørgen ArendtPublished in:9th European Conference on Synthetic Aperture RadarPublication date:2012Link back to DTU OrbitCitation (APA):Hemmsen, M. C., Hansen, J. M., & Jensen, J. A. (2012). Synthetic Aperture Sequential Beamformation appliedto medical imaging. In 9th European Conference on Synthetic Aperture Radar: EUSAR 2012 (pp. 34 - 37 ).IEEE.Synthetic Aperture Sequential Beamformation applied to medicalimaging.Martin Christian Hemmsen, Jens Munk Hansen, and Joergen Arendt JensenCenter for Fast Ultrasound Imaging, Department of Electrical Engineering,Technical University of Denmark, DK-2800 Kgs. Lyngby, DenmarkAbstractSynthetic Aperture Sequential Beamforming (SASB) is applied to medical ultrasound imaging using a multi elementconvex array transducer. The main motivation for SASB is to apply synthetic aperture techniques without the need forstoring RF-data for a number of elements and hereby devise a system with a reduced system complexity. Using a 192element, 3.5 MHz, λ-pitch transducer, it is demonstrated using tissue-phantom and wire-phantom measurements, how thespeckle size and the detail resolution is improved compared to conventional imaging.1 IntroductionIn multi-element Synthetic Aperture (SA) imaging, the ba-sic idea is to create a pressure wave from multiple ele-ments with a focused transmission. The concept of usingthe transmit focal point as a virtual source was introducedby Passmann and Ermert [1]. Virtual sources in SyntheticAperture Focusing (SAF) was further investigated by Fra-zier and O’Brien [2], Nikolov and Jensen [3, 4], and Baeand Jeong [5]. It was shown that the virtual source coin-cides with the focal point of the transducer, and that a depthindependent resolution can be achieved. Kortbek et al. in-troduced [6] the concept of Sequential Beamforming to SAimaging to reduce system requirements for real time imple-mentation. For a multi element linear array transducer, thelateral resolution can be made more range independent andimproved significantly compared to conventional imagingusing Dynamic Receive Focusing (DRF).The basic idea in Synthetic Aperture Sequential Beam-forming (SASB) is to create a dual-stage procedure usingtwo separate beamformers. In the initial stage a beam-former using a single focal point in both transmit and re-ceive, creates a set of focused scan lines. A second stagebeamformer creates a set of high resolution image pointsby combining information from multiple first stage focusedscan lines, which results in a dynamically expanding arrayas the image depth increases and a more range independentlateral resolution is obtained.This paper evaluates the feasibility of applying SASB tomedical ultrasound imaging using a multi element convexarray transducer. The method is assessed and performanceis quantized through a combination of water tank and tis-sue phantom measurements.2 MethodSASB creates a set of N focused scan lines from M emis-sions using a first stage beamformer with a fixed receivetime-delay profile. The delay profile is found from theround trip time-of-flight calculated astd(~rr) =1c(|~rtfp−~re|±2|~rfp−~rtfp|+ |~rr−~rtfp|). (1)The ± in (1) refer to whether the image point is above orbelow the transmit focal point. Figure 1 illustrates thetime-of-flight calculation for the SASB first stage beam-former.Figure 1: Wave propagation path (solid line) for calcu-lating the receive focusing time delays for a fixed focusedscan line.Each point in the focused scan line contains informationfrom a set of spatial positions limited by the opening angle,α = 2arctan(1/(2F#)), of the virtual source. A singleimage point is therefore potentially represented in multiplefirst stage focused scan lines.A second stage beamformer creates a set of high resolutionimage points by combining information from multiple firststage focused scan lines. A single sample, h, representingthe image point at location, ~rip, is calculated ash(~rip) =N∑k=1W(lθk , ~rip)lθk(tdk(~rip)), (2)where lθk(tdk(~rip)) is the sample at time tdk from the scanline, l, with propagation direction θk. The variable W isan apodization function with N values, which controls theweighting of the contribution from each of the N first stagescan lines. The time delay, tdk for the individual scan linesare found from the round trip time-of-flight,tdk(~rip) =2c(|~rV S − ~rθk | ± |~rip − ~rV S |), (3)where ~rθk is the scan line reference position, and ~rV S isthe position of the virtual source. Figure 2(a) illustratesthe time-of-flight calculation for the SASB second stagebeamformer.(a) Time-of-flight calculation (b) Apodization weight calculationFigure 2: Visualization of (a) time-of-flight calculationand (b) apodization weight calculation.The variable W in (2) is an apodization function determin-ing the weighting of the individual scan lines. The weightfor the scan line, lθk , contributing to a specific image pointcan be calculated from (4)W(lθk , ~rip) ={0.54 + 0.46cos(2nπ), if n is ≤ 10, if n is > 1(4)for the case of a desired Hamming apodization, where ncan be calculated fromn =d(lθk , ~rip)∆, (5)and is given as the ratio between the shortest distance,d(lθk , ~rip) from the image point to the scan line, and a de-sired aperture width, ∆. The distance d(lθk , ~rip) can beexpressed asd(lθk , ~rip) =‖(~rθk − ~rV Sk)× (~rV Sk − ~rip)‖‖~rθk − ~rV Sk‖(6)where ~rθk is the reference point of the scan line. Figure2(b) illustrates the apodization weight calculation.2.1 Performance IndicatorsDetail resolution The detail resolution is the separationat which identically point targets can be distinguished. It isdistinguished by the main lobe width of the point-spread-function (PSF) and the 6 dB axial pulse length, i.e. thelateral and axial full width at half maximum (FWHM) ofthe PSF.Cystic resolution The cystic resolution [7] is a mea-sure that considers the ability to detect anechoic or weaklyechogenic objects in the presence of strong surroundingobjects, and can be quantized by the clutter energy to totalenergy ratio (CTR). The CTR is defined as the ratio of theenergy outside a circular region R with radius R centeredat the peak of the PSF to the total PSF energyCTRR(~r0) = 10 log(1−∫~r∈R|h(~r, ~r0)|2dS∫|h(~r, ~r0)|2dS),where |h(~r, ~r0)| is the PSF at ~r0. To get a single measure,one either measures the drop in brightness for a fixed ra-dius R or the radius for which a cyst can be observed at afixed level, e.g. R12dB ≡ {R |CTRR = 12dB}.Tissue contrast resolution The theoretical model pro-posed by Smith et al. [8] is a measure that considers thetask of identifying a lesion within a uniform backgroundfor an ideal observer. The model states that the signal-to-noise ratio for lesion detection isSNR∆I =CdNeff1/2(SlatSax)1/2,where C is the contrast of a circular lesion of diameterd, Slat, and Sax, are, respectively, the lateral and axialspeckle dimension, and Neff is the number of independentimages. To improve detectability of a lesion with a givencontrast, it is therefore of great interest to maximize thedensityNeff(SlatSax), (7)which will be refered to as the speckle information density(SID). The number of independent images, Neff can be es-timated by measuring the signal-to-noise ratio at a point,SNR0 and using that the envelope-detected signals followRayleigh statistics [9]Neff =(SNR01.91)2. (8)The speckle dimensions, Slat and Sax can be found fromthe correlation cellSc =∫∞−∞CI(∆ξ)CI(0)d(∆ξ), (9)where CI is the spatial auto-covariance for the intensityand ∆ξ is the distance between image points.3 ResultsTo investigate the performance of SASB, a setup using acommercial available 3.5 MHz, 192 element, λ-pitch con-vex array transducer (Sound Technology Inc., State Col-lege, PA, USA) was used for water tank and tissue phan-tom measurements. The data acquisition was done usinga commercial available 2202 ProFocus ultrasound scanner(BK-Medical, Herlev, Denmark), capable of storing beam-formed baseband data [10]. For SASB and DRF imag-ing, a sliding 63 element sub-aperture was used for 384 fo-cused emissions respectively for SASB and DRF. Acqui-sitions were done interleaved and beamformed data werestored for off-line processing. For all emissions, a manu-facturer default excitation pulse was used with no transmitapodization on the active sub-aperture. The emission foreach sub-aperture was constructed using a fixed F-numbercorresponding to a focal depth of 69 mm. In receive SASBgenerated scan lines were created using a Gaussian receiveapodization (α = 0.5) applied to the same active sub-aperture as in transmit. DRF generated scan lines were cre-ated using a dynamic receive aperture maintaining a con-stant F-number of 0.8 and dynamic receive apodization us-ing a Gaussian apodization (α = 0.5).3.1 Water phantom measurementsTo evaluate the cystic resolution, and detail resolution awater tank phantom consisting of 4 wires at depths 16, 41,66, and 91 mm were used. The transducer was submergedinto the water and data were collected. Figure 3 illustratesthe measured PSF at depths of 41 and 91 mm as 6 dB con-tour plots with a dynamic range of 60 dB.Lateral position [mm]Axial position [mm]−6 −4 −2 0 2 4404244(a) PSF using DRF at 41 mmLateral position [mm]Axial position [mm]−6 −4 −2 0 2 4404244(b) PSF using SASB at 41 mmLateral position [mm]Axial position [mm]−6 −4 −2 0 2 4909294(c) PSF using DRF at 91 mmLateral position [mm]Axial position [mm]−6 −4 −2 0 2 4909294(d) PSF using SASB at 91 mmFigure 3: Measured PSF at depths of 41 and 91 mm ob-tained using a water tank phantom. The figures (a+c) ismade using DRF and (b+d) using SASB.Comparing the contours, the PSF using SASB is less rangedependent and has an improved lateral resolution in termsof the FWHM. The improvement in FWHM comes at acost in side-lobe level. The near side-lobes are higher us-ing SASB, they do however fall off faster than using DRF.Table 1: Contrast resolution, FWHM, and detail resolutionextracted from the water tank measurements.R12dB FWHMlat FWHMax Area6dB[mm] [mm] [mm] [mm2]DRF41mm 0.51 0.79 0.41 0.26SASB41mm 0.46 0.71 0.41 0.23DRF91mm 1.25 2.29 0.64 1.15SASB91mm 1.69 1.54 0.56 0.68From the results in Table 1 it follows that the axial andlateral resolution at FWHM is improved using SASB. Theaxial and lateral resolution at FWHM is respectively inaverage improved by 3% and 22%. The cystic resolutionat 12 dB is improved at 41 mm, but is degraded at 91 mm.The degradation at 91 mm is due to the high near side-lobelevel. Figure 4 illustrates the CTR together with the clutterlevel for the two wires. Note that the CTR for the SASBimages are in general equally good or better for any radius,but for radius between 1 and 4 mm. The clutter level isimproved using SASB for any radius.0 2 4 6−80−60−40−200Radius [mm]CTR [dB]  DRFSASB(a) CTR at 41 mm0 2 4 6 8 10−80−60−40−200Radius [mm]CTR [dB]  DRFSASB(b) CTR at 91 mmFigure 4: CTR for dynamic receive focusing (thicksolid lines) and synthetic aperture sequential beamform-ing (thick dashed lines). Clutter levels are shown below asthin dashed and solid lines.3.2 Tissue phantom measurementsTo investigate the performance for lesion detection, a tis-sue mimicking phantom was scanned and speckle size, andSNR0 were calculated for 20x20 mm regions at depthsranging from 20 to 140 mm. Table 2, shows the resultfor the two regions with centers at 41 and 91 mm. Fromthe figure it follows that the speckle size is on average re-duced by 15.3%, and the speckle information density isimproved by 12% using SASB. Figure 5(a) illustrates thespeckle size and speckle information density in the rangefrom 20 to 140 mm. Figure 6 illustrates two regions fromthe tissue phantom with small anechoic objects for a visualcomparison of the tissue contrast resolution.Table 2: Signal-to-noise ratio at a point, SNR0, specklesize and the density Neff/(SaxSlat).SNR0 Speckle size Neff/(SaxSlat)[mm2] [mm-2]DRF41mm 1.87 0.32 2.37SASB41mm 1.86 0.27 2.73DRF91mm 1.85 0.54 1.37SASB91mm 1.87 0.46 1.6320 40 60 80 100 120 1401234Depth [mm]Neff/(SaxS lat )  DRFSASB(a) Speckle Information Density20 40 60 80 100 120 1400.20.40.60.8Depth [mm]Speckle Size [mm2]  DRFSASB(b) Speckle sizeFigure 5: Visualization of (a) speckle information densityand (b) speckle size calculated from the tissue phantommeasurement, using SASB (dashed line) and DRF (solidline).Lateral position [mm]Axial position [mm]−10 0 10303540455055(a) DRFLateral position [mm]Axial position [mm]−10 0 10303540455055(b) SASBLateral position [mm]Axial position [mm]−10 0 1080859095100(c) DRFLateral position [mm]Axial position [mm]−10 0 1080859095100(d) SASBFigure 6: B-mode images of the tissue phantom with smallanechoic objects (a+c) using DRF and (b+d) using SASB.4 ConclusionSASB has been successfully implemented for a convexmulti element array transducer. Evaluation using water andtissue phantom measurements indicate a performance im-provement for detail resolution, speckle size, and speckleinformation density. The evaluation show that SASB isable to increase the image range of a uniform lateral res-olution and to obtain a more isotropic point spread func-tion. The results presented in this paper proves the via-bility of the method for real-time 2D ultrasound imagingusing commercially available equipment.References[1] C. Passmann and H. Ermert, A 100-MHz ultrasoundimaging system for dermatologic and ophthalmologicdiagnostics, IEEE Trans. Ultrason., Ferroelec., Freq.Contr., vol. 43, pp. 545–552, 1996[2] C. H. Frazier and W. D. O’Brien, Synthetic aper-ture techniques with a virtual source element, IEEETrans. Ultrason., Ferroelec., Freq. Contr., vol. 45, pp.196–207, 1998[3] S. I. Nikolov and J. A. Jensen, Virtual ultrasoundsources in high-resolution ultrasound imaging, inProc. SPIE - Progress in biomedical optics and imag-ing, vol. 3, pp. 395–405, 2002[4] ——, 3D synthetic aperture imaging using a virtualsource element in the elevation plane, in Proc. IEEEUltrason. Symp., vol. 2, pp. 1743–1747, 2000[5] M. H. Bae and M. K. Jeong, A study of synthetic-aperture imaging with virtual source elements in B-mode ultrasound imaging systems, in IEEE Trans.Ultrason., Ferroelec., Freq. Contr., vol. 47, pp. 1510–1519, 2000[6] J. Kortbek, J. A. Jensen, and K. L. Gammelmark,Synthetic aperture sequential beamforming, Ultra-sonics, p. Submitted, 2009[7] D. Vilkomerson, J. Greenleaf, and V. Dutt, Towardsa Resolution Metric for Medical Ultrasound Imaging,in Proc. IEEE Ultrason. Symp., pp. 1405–1410, 1995[8] S. W. Smith, R. F. Wagner, J. M. Sandrik, andH. Lopez, Low contrast detectability and con-trast/detail analysis in medical ultrasound, Ultrason.Imaging, vol. 30, no. 3, pp. 164–173, May 1983[9] R. F. Wagner, S. W. Smith, J. M. Sandrick, andH. Lopez, Statistics of speckle in ultrasound B-scans,IEEE Trans. Son. Ultrason., vol. 30, pp. 156–163,1983[10] M. C. Hemmsen, S. I. Nikolov, M. M. Pedersen,M. J. Pihl, M. S. Enevoldsen, J. M. Hansen, and J. A.Jensen, Implementation of a versatile research dataacquisition system using a commercially availablemedical ultrasound scanner, IEEE Trans. Ultrason.,Ferroelec., Freq. Contr., p. Accepted, 2011

Synthetic Aperture Sequential Beamformation applied to medical imaging.

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Synthetic Aperture Sequential Beamformation applied to medical imaging.

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