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199Chapter seven: Contrast media and harmonic imaging
Renal tumor
Perfusion Time (s)
y
OpticalImageofaBurstingBubble
34
27
20
13
7
0
Figure 7.18 Rell time of a kidney estimated by quantifying echogenicity of the
organ following the injection of a contrast agent. The center of the kidney has a
tumor. Its rell time is very different from those of the surrounding structures.
(Courtesy of Prof. Kathy Ferrara at University of California at Davis.)
passively to an organ of interest, or it may actively seek out the targeted
tissue by binding it with specic molecules that interact only with a certain group of molecules in a tissue (Lanza et al., 2000; Ferrara et al., 2007).
Figure7.19 shows a sequence of optical images on the left acquired by a
high-speed camera as a bubble bursts, maybe induced by insonication.
The electron micrograph on the right shows a bubble with an oil shell that
0.5 µm
5 µm
EMimageofdrugdeliver
vehiclewiththickoillayer
designedtocarrydrugs
0 0.4 0.8 1.2
1.6 2.0 2.4
Time (µsec)
Figure 7.19 Snapshots of the rupture of a bubble at different times, photographed by a high-speed camera. (Courtesy of Prof. Kathy Ferrara at University
of California at Davis.)
200 Diagnostic ultrasound: imaging and blood ow measurements
can be loaded with a drug to be released upon ultrasound insonication.
In this case, the bubbles are passively carried by the blood stream to the site
of interest. The encapsulated bubble may also be conjugated with a targeting ligand, which targets a certain molecule on the endothelial surface so
that the bubble actively seeks the molecule and binds with it (Figure7.15).
The drug that the bubble carries is then released upon insonication.
References and Further Reading Materials
Averukiou MA. Tissue harmonic imaging. IEEE Ultrasonics Symp Proc 2000; 2:
1563–1572.
Chang PP, Shung KK. Interaction of ultrasound with contrast agents. In Thomsen
HH, Muller RN, and Mattery RF (eds.), Trends in contrast agents. Berlin:
Springer, 1998.
Crum LA and Prosperetti A. Nonlinear oscillations of gas bubbles in liquids:
An interpretation of some experimental results. J Acoust Soc Am 1983; 73:
121–127.
de Jong N. Improvements in ultrasound contrast agents. Eng Med Biol Mag 1996;
15: 72–82.
de Jong N, Cornet R, and Lancee CT. High harmonics of vibrating gas-lled micro-
spheres. Part I. Simulations. Ultrasonics 1994; 32: 447–453.
Ferrara K, Pollard R, and Borden M. Ultrasound microbubble contrast agents:
Fundamentals and application to gene and drug delivery. Annu Rev Biomed
Eng 2007; 9: 415–447.
Gessner R and Dayton P. Advances in molecular imaging. Mol Imaging 2010; 9:
117–127.
Goldberg BB, Liu JB, and Forsberg F. Ultrasound contrast agents: A review.
Ultrasound Med Biol 1994; 20: 319–333.
Lanza G, Hall C, Scott, M, Fuhrhop R, March J, and Wickline S. Molecular imaging
with targeted ultrasound contrast agent. IEEE Ultrasonics Symp Proc 2000; 2:
1917–1926.
Leighton TG. The acoustic bubble. San Diego: Academic Press, 1994.
Medwin H. Counting bubbles acoustically: A review. Ultrasonics 1977; 15: 7–13.
Morse PM and Ingard KU. Theoretical acoustics. New York: McGraw Hill, 1968.
Plesset MS. The dynamics of cavitation bubbles. J Appl Mech 1949; 16: 277–282.
Rayleigh L. On the pressure developed in a liquid during the collapse of a spherical
cavity. Phil Mag 1917; 34: 94–98.
Shi W and Forsberg F. Ultrasonic characterization of the nonlinear properties of
contrast microbubbles. Ultrasound Med Biol 2000; 26: 93–104.
Wilson SR and Burns P. Microbubble-enhanced US in body imaging: What role?
Radiology 2010; 257: 24–39.
chapter 8
Intracavity and
high-frequency (HF) imaging
Conventional ultrasonic imaging systems typically use frequencies from
2 to 15 MHz. Scanners at lower frequencies have the advantage of a larger
depth of penetration but suffer from poorer resolution. To improve spatial resolution, one obvious strategy would be to increase the frequency.
The axial resolution is determined by the pulse duration or bandwidth
of the pulse. For a xed number of cycles per pulse, an increase in frequency would result in a reduction in wavelength and thus pulse duration. The relationship between frequency (wavelength) and lateral spatial
resolution is given by Equation (3.29). These relationships are graphically
illustrated in Figures8.1 and 8.2. As ultrasound frequency is increased to
50 MHz, and an axial resolution and lateral resolution of better than 20
and 100 μm for an f# of 2.9 can be achieved, respectively. The price to be
paid is an increase in attenuation. The effect of attenuation coefcient of a
few types of tissues of clinical interest is shown in Figure8.3. At 50 MHz,
the depth of penetration for most tissues would be limited to 4–5 mm.
Although, as was discussed, the depth of penetration may be increased
slightly by introducing novel signal processing methods such as coded
excitation, the range of frequencies that can be applied to a certain organ
is limited.
8.1 Intracavity imaging
Intracavity imaging such as transesophageal, transrectal, and transvaginal imaging is a partial solution to achieving improvements in spatial resolution. Since imaging organs like the heart, prostate, and uterus/ovary
from the body surface does not usually allow the utilization of frequencies higher than 5 MHz because they are deep-lying organs, probes may
be modied to be inserted through open cavities of the body to be placed
closer to these organs to allow higher frequencies to be used.
8.1.1 Transesophageal cardiac imaging
Phased arrays consisting of more than 48 elements at frequencies from 5
to 7.5 MHz can be mounted on the tip of an endoscope with a diameter
201
202 Diagnostic ultrasound: imaging and blood ow measurements
Axial Resolution vs. Bandwidth
100
80
60
(µm)
40
axial
R
20
0
0255075 100 125 150
Bandwidth (MHz)
Figure 8.1 Calculated axial resolution as a function of bandwidth.
less than 10 mm for imaging the heart from the esophagus. The endoscope allows the manipulation of the position and direction of the phased
array. During scanning, the transesophageal probe is inserted into the
esophagus and the tip is positioned against the wall of the esophagus
under local anesthesia. A majority of the probes are capable of biplane
imaging; i.e., two orthogonal images are produced. More advanced versions can produce images in any direction by mechanically rotating the
array. The most advanced probes have 2D arrays, allowing 3D imaging
Laterial Resolution vs.
300
250
200
(µm)
150
laterial
R
100
50
0
0255075 100 125 150
Center Frequency
f# = 2.9
f# = 0.7
Center Frequency f0 (MHz)
Figure 8.2 Calculated lateral resolution as a function of ultrasound center
freque nc y.
203Chapter 8: Intracavity and high-frequency (HF) imaging
Depth (mm)
At 50 MHz
α
α
α
α
≈ 2.5 dB/mm
blood
≈ 1.1 dB/mm
cornea
≈ 1.7 dB/mm
iris
≈ 10 dB/mm
skin
0.9
0.8
0.7
0.6
0.5
0.4
0.3
Trasnfer Function
0.2
0.1
Transfer Function vs. Depth at
1
0
0 0.5 1 1.5 2
50 MHz
Cornea
Skin section
Blood
Iris
Figure 8.3 The effect of ultrasound attenuation of four different types of tissues
plotted as a function of depth at 50 MHz. The transfer function is given by e
–βz
where β is the attenuation coefcient and z is the propagation depth.
of the heart in real time, which will be discussed in Chapter 8 in more
detail. Transesophageal imaging of the heart yields better images of
not only the whole heart because of the higher frequencies, but also the
base of the heart, which cannot be adequately accessed by transthoracic
imaging. An additional benet is that transesophageal imaging allows
continuous monitoring of the cardiac functions, which has proven valuable in anesthesiology during surgery. A photo of such a probe is shown
in Figure8.4. Commercial catheters (10 French units, 1 F = 0.33 mm outer
diameter) mounted near the tip on the side of a linear array at a frequency
of 8 MHz are also available for intracardiac imaging. The catheter can be
guided to the heart with a guide wire via a peripheral artery.
,
8.1.2 Transrectal and transvaginal imaging
Probes at frequencies higher than 5 MHz are available for most imaging systems for insertion into the rectum or vagina for better imaging
of the prostate and uterus/ovary. A linear array or curved linear array is
mounted on the side or at the tip of a probe. A full bladder, which used to
be recommended for transabdominal obstetrical imaging of a fetus, may
now be replaced by transvaginal imaging. A photo of several transrectal
and transvaginal probes is shown in Figure8.5.
204 Diagnostic ultrasound: imaging and blood ow measurements
Control housing
Flexible shaft
ew
Enlarged vi
of the tip
Figure 8.4 A photo of a transesophageal probe. (Courtesy of Oldelft B.V., Delft,
The Netherlands.)
8.1.3 Endoluminal imaging
Catheter-based imaging systems have also been used to image the gastrointestinal tract, including the colon, esophagus, and stomach (Liu and
Goldberg, 1999). A few manufacturers have developed specialized ultrasonic imaging systems to accomplish the same by mounting ultrasonic
transducers/arrays on the end of an endoscope. The frequency of the
ultrasound probe ranges from 7 to 20 MHz, with a uid-lled balloon at
the tip. Linear arrays and radial arrays (elements mounted on the circumference of an endoscope) have been introduced recently.
Figure 8.5 A photo showing several types of ultrasound probes, including linear
arrays, linear curved arrays, Doppler probes, and transrectal and transvaginal
probes (long and slender ones in the upper half of the photo). (Courtesy of Sound
Technology, Inc., State College, PA.)
205Chapter 8: Intracavity and high-frequency (HF) imaging
8.2 Intravascular imaging
Imaging of the wall of blood vessels for the purpose of estimation of
the degree of stenosis and characterization of atherosclerotic plaques
has been pursued for many years with a variety of imaging modalities
(Pandian, 1989; Liu and Goldberg, 1999). X-ray angiography has been
the gold standard in the past for assessing stenosis. The drawbacks of
x-ray are that (1) it is a form of ionizing radiation involving the injection
of a contrast agent, and (2) it is a 2D projection image of a 3D structure. More than two views are necessary to have a more accurate assessment of the stenotic vessel. As a result, its role is being challenged by
both magnetic resonance imaging and ultrasound (Shung et al., 1992).
Plaque composition characterization is of clinical importance in that it
has been hypothesized that vulnerable plaques consisting of a lipid core
with a brous cap are most likely to rupture, causing the formation of
clots and serious clinical consequences, such as stroke and heart attack.
Imaging options for characterizing plaque composition are quite limited.
Fiberoptic angioscopy, in which an optic ber is introduced via catheterization to the site of interest for the visualization of plaque surface, and
transcutaneous ultrasound have been used. The former procedure, which
involves injection of saline for ushing out light opaque blood, can only
visualize the lesion surface, whereas the latter suffers from poor resolution. Intravascular ultrasound and optical coherent tomography (OCT)
are possible alternatives to alleviate these problems (Liu and Goldberg,
1999; Bouma and Tearney, 2002). Intravascular ultrasound scanners typically are operated in the frequency range from 20 to 60 MHz, depending
upon the imaging catheter used. There are two different types of imaging catheters on the market today. In one, a single-element transducer at
a frequency from 30 to 60 MHz of 1.75 mm diameter, making an angle
of 10° relative to the direction normal to the long axis of the catheter,
is mounted near the tip of the catheter. The transducer is mechanically
rotated at a very high speed (~1800 rpm) by a shaft. These catheters have
a size of 3.5 to 6 F. In another, a 64-element array at 20 MHz having a
1.5λ pitch is mounted around the circumference of a 3.5 F catheter (1.2
mm outer diameter). The elevational width of the array is 0.7 mm. Also
mounted on the catheter with the array are several integrated circuit
chips that perform the functions of low-noise broadband preamplication and multiplexing. A synthetic imaging approach in which 1 element
transmits and 14 elements receive is used to form the image. Figure8.6
shows a catheter with a mechanically rotated transducer (top) and a
catheter with an array wrapped around the circumference (bottom). The
array catheter, although it has a higher frame rate, yields an image quality slightly inferior to that of the mechanically rotated type, primarily
because of its lower frequency.
206 Diagnostic ultrasound: imaging and blood ow measurements
Ca
od
• Mechanical Transducer
•
(b)
–Transducer that rotates on a drive shaft
(a)
Array/Solid State Transducer
Figure 8.6 (a) An intravascular imaging catheter with a mechanically rotated
transducer. (b) A catheter with an array wrapped around the circumference.
(Courtesy of Boston Scientic and Volcano Therapeutics.)
In addition to plaque characterization, intravascular ultrasound has
been found to be useful in guiding the placement of a stent and monitoring stent restenosis. Figure 8.7 shows a 40 MHz image of an artery,
whereas Figure8.8 shows a 20 MHz image of a stent. At present, all intravascular ultrasonic imaging devices use side-looking catheters, which
lack the capability of visualizing anatomic structures in front of the catheter. This capability is important in not causing any injury to the blood
vessel itself, while being guided to the site of interest, such as perforation
and dissection. Efforts are now underway in developing forward-looking
Figure 8.7 An image of an artery acquired at 40 MHz by a catheter with a rotating
single-element transducer. (Courtesy of Boston Scientic.)
Plaque
Blo
theter
207Chapter 8: Intracavity and high-frequency (HF) imaging
Figure 8.8 A 20 MHz imaging catheter is placed at the mid-stent level. The right
panel shows a longitudinal view of the stent. The echogenic stent structure is
clearly seen. (Courtesy of Volcano Therapeutics.)
intravascular catheters (Degertekin et al., 2006), and a forward-looking
device that is a variation of the rotating single-element type is now commercially available.
A more recent trend in intravascular imaging is in combining two
or more modalities taking advantage of the merits of each (Li et al., 2010).
This will be discussed in more detail in Section 9.4.2.
8.3 High-frequency imaging
Scanners operated at frequencies higher than 20 MHz have been developed for applications in ophthalmology, dermatology, and small animal
imaging. Typically these devices, called ultrasonic backscatter microscope or ultrasonic biomicroscope (UBM), obtain images by scanning a
single-element ultrasonic transducer either in a sector format or linearly.
The construction of a UBM is identical to that of a static B-mode scanner
(Briggs and Arnold, 1996; Pavlin and Foster, 1995). Scanning can also be
achieved by better utilizing the focus of the transducer by incrementally
moving the transducer in the axial direction, called B-D (D stands for
208 Diagnostic ultrasound: imaging and blood ow measurements
Figure 8.9 A 256-element HF 30 MHz linear array of one-wavelength pitch fabricated with conventional dicing technology.
depth) mode scan. A composite image is formed by combing the focused
segments of multiple images acquired as the transducer is moved in the
axial direction. This mode of scanning improves lateral resolution by sacricing frame rate. Commercial UBMs can achieve a frame rate of 30 per
second because the excursion range is extremely small. High-frequency
scanners that utilize linear arrays have been commercially available for
preclinical small imaging since 2009 (Foster et al., 2009). Extensive investigations are still being undertaken in developing HF arrays and associated
imaging electronics (Ritter et al., 2002; Foster et al., 2009; Cannata et al.,
2011; Hu et al., 2011) for preclinical and clinical applications. Conventional
dicing technology may be used to fabricate linear arrays up to 50 MHz
because dicing blades are only capable of dicing kerfs as small as 15 μm.
For arrays of higher than 50 MHz, alternative technology such as micromachining, which will be discussed in Chapter 9, laser dicing (Foster
et al., 2009), or deep reactive ion etching (DRIE) (Yuan et al., 2008) may
have to be exploited. A 256-element HF 30 MHz linear array of one-wavelength pitch fabricated with conventional dicing technology is shown in
Figure8.9. The array with an aperture size of 2 × 6 mm had a bandwidth
of 55% and a cross talk level of –27 dB. The design philosophy is similar to
low-frequency arrays with one or more matching layers and a light backing. A laser- (excimer at 248 nm) diced 256-element 30 MHz array with
a kerf width of 8 μm for a preclinical scanner (VEVO 2100) produced by
Visualsonics (Toronto, Canada) is shown in Figure8.10. The corresponding images obtained by this array, shown in Figure8.11, clearly demonstrate the superior quality of the image, where all voids in the phantom
are visualized, over that obtained with a UBM. This scanner is capable of
performing conventional and color Doppler.
There are many clinical applications for high-frequency ultrasound.
In ophthalmology, scanners at 20 MHz or slightly lower have been used
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