**2. Recent advances in ultrasound imaging**

Sound in the human audible range have frequencies between 20 and 20 000 Hz. Ultrasound is defined as sound with frequencies above 20 kHz. In medical imaging, the ultrasound frequen‐ cies are usually between 2 and 40 MHz, with the highest frequencies currently used in intravascular ultrasound (IVUS).

The generation of an ultrasound image is based on transmission of sound pulses and receiving the echoes that have been reflected from tissue boundaries or scattered from smaller objects. In most conventional scanners today, a narrow ultrasound beam is transmitted from the ultrasound transducer. When the transmitted pressure pulse meets a hinder in form of a boundary between different soft tissues, or scatter points within the tissue with different acoustic properties, some of the energy of the transmitted sound pulse is echoed back to the transducer. This pulse-echo principle forms the basis of all ultrasound-imaging techniques, such as conventional brightness mode (B-mode) imaging of organs, imaging of blood flow using Doppler techniques and exploration of mechanical tissue properties using ultrasound elastography techniques.

**2.2. Ultrasound elastography**

**Figure 3.** Pocket-sized ultrasound (Vscan from GE Healthcare)

with differences in strain (figure 4).

tumour as displayed on an Ultrasonix MDP scanner.

The concept of ultrasound imaging of tissue strain or elasticity is often referred to as ultrasound elastography and the corresponding 2D images are frequently called elastograms. The imaging technique is often explained to be analogue to palpation, where the physician uses the fingers to apply a slight pressure in order to examine the stiffness of the tissue. If a organ is vibrating or excited, ultrasound elastography methods can in a similar fashion be used to map areas

Ultrasound-Based Guidance and Therapy http://dx.doi.org/10.5772/55884 31

**Figure 4.** Elastography. A) Ultrasound B-mode image of a small meningioma, and B) the ultrasound elastogram of the

The theoretical framework for the study of behavior of vibrating soft tissue was establish‐ ed in the early 1950ies. Von Gierke et al. published *"Physics of vibrations in living tissues*" in 1952 [5], for example. However, it was not until 30 years later that tissue movement was first measured for clinical purposes by using ultrasound in a study of tissue motion in the liver caused by vascular pulsation [6, 7]. In the late 1980ies, techniques for vibration elastogra‐ phy imaging, also known as vibration amplitude sonoelastography or simply sonoelastici‐ ty imaging was developed [8]. In this technique a low frequency vibration (20-1000 Hz) is applied externally to the skin surface to investigate the subcutaneous structures. The internal

#### **2.1. Advances in ultrasound hardware and transducer technology**

The ultrasound machines and ultrasound probes have gone through massive improvements in the last decade. The general increase in computer power is opening new possibilities for implementing sophisticated methods for beam forming. This may lead to higher resolution and better image quality than for existing solutions [1]. The general trend with miniaturization of components has also strongly influenced the size of the ultrasound imaging systems. Small handheld ultrasound devices have been developed, which makes ultrasound an extremely portable imaging technology. One example of such a pocket sized ultrasound device is the Vscan from GE Healthcare (figure 3), which has been explored for use in echocardiography [2]. The ultrasound transducer technology has made tremendous progress the last decade. The number of elements used by a transducer is increasing and the trend is to go from a single row of elements (1D) to multi-row arrays (1.25D / 1.5D) and 2D matrix arrays. The latter provides the possibility to perform 4D ultrasound imaging, in which a 3D ultrasound volume is acquired and displayed in real time. 4D ultrasound imaging may also be used for monitoring of treatment, e.g. radiofrequency ablation [3].

Ultrasound arrays today are mostly based on piezoelectric materials. The research activities in MUT (Micromachined Ultrasound Transducer) technology, and perhaps especially CMUT (capacitive MUT) transducers, pave the way for silicon-based arrays [4]. This may introduce probes that are cheaper, more customizable and have higher frequencies and bandwidth compared to piezoelectric transducers. In combination with the everlasting trend of miniatur‐ ization, the CMUTs may in a long-term perspective allow complete ultrasound systems to be seamlessly integrated with surgical tools. It may very well be that the future surgical instru‐ ment has an ultrasound transducer integrated on the tip, and a display unit integrated in the handle.

**Figure 3.** Pocket-sized ultrasound (Vscan from GE Healthcare)

#### **2.2. Ultrasound elastography**

**2. Recent advances in ultrasound imaging**

30 Advancements and Breakthroughs in Ultrasound Imaging

**2.1. Advances in ultrasound hardware and transducer technology**

intravascular ultrasound (IVUS).

elastography techniques.

treatment, e.g. radiofrequency ablation [3].

handle.

Sound in the human audible range have frequencies between 20 and 20 000 Hz. Ultrasound is defined as sound with frequencies above 20 kHz. In medical imaging, the ultrasound frequen‐ cies are usually between 2 and 40 MHz, with the highest frequencies currently used in

The generation of an ultrasound image is based on transmission of sound pulses and receiving the echoes that have been reflected from tissue boundaries or scattered from smaller objects. In most conventional scanners today, a narrow ultrasound beam is transmitted from the ultrasound transducer. When the transmitted pressure pulse meets a hinder in form of a boundary between different soft tissues, or scatter points within the tissue with different acoustic properties, some of the energy of the transmitted sound pulse is echoed back to the transducer. This pulse-echo principle forms the basis of all ultrasound-imaging techniques, such as conventional brightness mode (B-mode) imaging of organs, imaging of blood flow using Doppler techniques and exploration of mechanical tissue properties using ultrasound

The ultrasound machines and ultrasound probes have gone through massive improvements in the last decade. The general increase in computer power is opening new possibilities for implementing sophisticated methods for beam forming. This may lead to higher resolution and better image quality than for existing solutions [1]. The general trend with miniaturization of components has also strongly influenced the size of the ultrasound imaging systems. Small handheld ultrasound devices have been developed, which makes ultrasound an extremely portable imaging technology. One example of such a pocket sized ultrasound device is the Vscan from GE Healthcare (figure 3), which has been explored for use in echocardiography [2]. The ultrasound transducer technology has made tremendous progress the last decade. The number of elements used by a transducer is increasing and the trend is to go from a single row of elements (1D) to multi-row arrays (1.25D / 1.5D) and 2D matrix arrays. The latter provides the possibility to perform 4D ultrasound imaging, in which a 3D ultrasound volume is acquired and displayed in real time. 4D ultrasound imaging may also be used for monitoring of

Ultrasound arrays today are mostly based on piezoelectric materials. The research activities in MUT (Micromachined Ultrasound Transducer) technology, and perhaps especially CMUT (capacitive MUT) transducers, pave the way for silicon-based arrays [4]. This may introduce probes that are cheaper, more customizable and have higher frequencies and bandwidth compared to piezoelectric transducers. In combination with the everlasting trend of miniatur‐ ization, the CMUTs may in a long-term perspective allow complete ultrasound systems to be seamlessly integrated with surgical tools. It may very well be that the future surgical instru‐ ment has an ultrasound transducer integrated on the tip, and a display unit integrated in the The concept of ultrasound imaging of tissue strain or elasticity is often referred to as ultrasound elastography and the corresponding 2D images are frequently called elastograms. The imaging technique is often explained to be analogue to palpation, where the physician uses the fingers to apply a slight pressure in order to examine the stiffness of the tissue. If a organ is vibrating or excited, ultrasound elastography methods can in a similar fashion be used to map areas with differences in strain (figure 4).

**Figure 4.** Elastography. A) Ultrasound B-mode image of a small meningioma, and B) the ultrasound elastogram of the tumour as displayed on an Ultrasonix MDP scanner.

The theoretical framework for the study of behavior of vibrating soft tissue was establish‐ ed in the early 1950ies. Von Gierke et al. published *"Physics of vibrations in living tissues*" in 1952 [5], for example. However, it was not until 30 years later that tissue movement was first measured for clinical purposes by using ultrasound in a study of tissue motion in the liver caused by vascular pulsation [6, 7]. In the late 1980ies, techniques for vibration elastogra‐ phy imaging, also known as vibration amplitude sonoelastography or simply sonoelastici‐ ty imaging was developed [8]. In this technique a low frequency vibration (20-1000 Hz) is applied externally to the skin surface to investigate the subcutaneous structures. The internal

motion of the tissue is investigated with a pulsed Doppler technique. Stiff tissue responds differently to the vibrations than softer tissue, and can therefore be distinguished in the realtime images.

**2.3. Nonlinear acoustics and contrast agents**

multiple scattering [28-31].

reflector enhancing the problem with multiple scattering.

In 1980, Carstensen and Muir published two papers describing the importance of nonlinear acoustics within the field of medical ultrasound imaging [26, 27]. These papers predicted and demonstrated nonlinear acoustical effects relevant for intensities and frequencies common in biomedical imaging. There has been an increasing interest with respect to nonlinear biomedical acoustics during the last 30 years. This interest was further escalated by the introduction of ultrasound contrast agents in the form of microbubbles and the study of these microbubbles

Ultrasound-Based Guidance and Therapy http://dx.doi.org/10.5772/55884 33

Nonlinear effects can be important in the forward wave propagation. The back-scattered pressure levels of the echoes are typically too low to induce any significant nonlinear effects. One source of nonlinear terms is produced by the deformation of tissue volume elements during compression and expansion with strongly curved phase fronts. It is, however, common to use transmit beams with relatively smooth phase fronts. Consequently, this nonlinear source is usually not the most dominant. The other important nonlinear source is nonlinear terms in the tissue elasticity and hence in the relation between acoustic pressure and tissue compres‐ sion/expansion. Nonlinear terms in the tissue elasticity are responsible for the fact that the tissue becomes stiffer during compression and softer during expansion. The compression also increases the mass density of the tissue, but this effect is inferior to the increased stiffness and the propagation velocity and will therefore be pressure dependent and will increase with increasing compressions and thus with increasing pressure. The resulting distortion of the transmit pressure field produces harmonic components which today are utilized in tissue harmonic imaging, especially in transcutaneous cardiac and abdominal imaging to suppress

Ultrasound imaging is based on several assumptions, and one important assumption is that multiple scattering is neglected. For many organs, this approximation is valid. However, for the body wall, where larger variations in material parameters often are found, this assumption can be inadequate. Interfaces between soft tissue components with significant differences in material parameters give so strong echoes from the transmitted acoustic pulses that multiple scattering can get significant amplitudes. Such multiple scatterings are usually termed pulse reverberations [32, 33]. These reverberations reduce the ratio of the strongest to the weakest scatterer that can be detected in the neighborhood of each other, defined as the contrast resolution in the image. Reduced contrast resolution is in particular a problem when imaging hypo-echoic structures such as the heart chambers, the lumen of large blood vessels, some atherosclerotic lesions, cysts, some tumors, the gallbladder as well as in fetal imaging. The contact interface between the ultrasound transducer itself and the soft tissue is also a strong

Ultrasound contrast agents are made as a suspension of gas microbubbles encapsulated in thin stabilizing shells made from lipid or albumin. Typical bubble size is in the 1-5 μm range and the contrast bubbles are intravenously injected to increase the scattering from blood, which is weak compared to the scattering from soft tissues. Commercially available contrast bubbles are stable and small enough to enable transpulmonary passage and the blood half-life is typically in the range of 1-10 minutes. Scattering from microbubbles occurring within a liquid

was the main impetus for the introduction of the tissue harmonic imaging technique.

In the early 1990ies, the development of compression elastography, also referred to as quasistatic elasticity imaging, begun. Ophir published a paper in 1991 where ultrasound radio frequency (RF) data before and after applying compression were compared and processed using cross-correlation to obtain the time-shifts of the echoes. This allowed the subsequent calculation of elastograms [9]. The quasi-static elasticity imply that the force is applied for a sufficiently long time for the tissue strain to stabilize, and the resulting difference in echo travel time between ultrasound data acquired before and after compression can be calculated. The tissue may also be excited by applying forces at the surface (manually or by electromechanical devices) or by physiological processes within the organ, as for example the pulsation of the arteries. The generated elastograms are usually displayed as a color-coded overlay on the conventional ultrasound brightness mode image. The color mapping may cover a range of unit-less strain values as percentages from minimum (negative) strain to maximum (positive) strain. Alternatively, it may also be mapped from "soft" to "hard" tissue, thereby not quanti‐ fying the strain range displayed. Quasi-static elasticity imaging has been evaluated in a broad range of clinical applications. It has been reported used in diagnostics of tumors in for example breast, prostate, liver, the thyroid gland and in the brain (figure 4) [10-15]. Quasi-static elasticity imaging is an emerging ultrasound imaging modality, now becoming more and more available as an option on commercial ultrasound systems.

As previously explained, the elastography methods require that the tissue is excited. The tissue movement can be caused by physiological processes internally in the organ such as the pulsation of the arteries. The tissue can also be externally excited by manually pushing the tissue or by using an electromechanical vibrating device. An alternative approach is to use the acoustic radiation force of an ultrasonic focused beam to generate displacements in the tissue with subsequent detection of the mechanical properties. One example of such an approach is the Acoustic Radiation Force Impulse (ARFI) method developed at Duke University [16]. In this technique, short duration acoustic pulses (push pulses) are used to generate small localized displacements deep in the tissue. These displacements are tracked by ultrasonic cross corre‐ lation, in a similar fashion as for the quasi-static elasticity imaging. The method has been investigated for imaging of focal liver lesions, prostate and breast [17-19].

Another example is the innovative Supersonic Shear Imaging (SSI) method developed by the research group at the Laboratoire Ondes et Acoustique [20]. In SSI the acoustic radiation force is used to generate low-frequency shear waves (50-500 Hz) remotely in the tissue. The shear modulus of the tissue can be quantified by imaging the share wave propagation in the tissue by using ultrasound frame rates of several kHz. The method has been explored for diagnosis of liver fibrosis, breast lesions and cornea [21-23].

For a more detailed overview about methods for ultrasound elasticity imaging and its clinical use we recommend to read the review papers by Wells and Liang [24] and Parker, Doyley and Rubens [25].

#### **2.3. Nonlinear acoustics and contrast agents**

motion of the tissue is investigated with a pulsed Doppler technique. Stiff tissue responds differently to the vibrations than softer tissue, and can therefore be distinguished in the real-

In the early 1990ies, the development of compression elastography, also referred to as quasistatic elasticity imaging, begun. Ophir published a paper in 1991 where ultrasound radio frequency (RF) data before and after applying compression were compared and processed using cross-correlation to obtain the time-shifts of the echoes. This allowed the subsequent calculation of elastograms [9]. The quasi-static elasticity imply that the force is applied for a sufficiently long time for the tissue strain to stabilize, and the resulting difference in echo travel time between ultrasound data acquired before and after compression can be calculated. The tissue may also be excited by applying forces at the surface (manually or by electromechanical devices) or by physiological processes within the organ, as for example the pulsation of the arteries. The generated elastograms are usually displayed as a color-coded overlay on the conventional ultrasound brightness mode image. The color mapping may cover a range of unit-less strain values as percentages from minimum (negative) strain to maximum (positive) strain. Alternatively, it may also be mapped from "soft" to "hard" tissue, thereby not quanti‐ fying the strain range displayed. Quasi-static elasticity imaging has been evaluated in a broad range of clinical applications. It has been reported used in diagnostics of tumors in for example breast, prostate, liver, the thyroid gland and in the brain (figure 4) [10-15]. Quasi-static elasticity imaging is an emerging ultrasound imaging modality, now becoming more and more available

As previously explained, the elastography methods require that the tissue is excited. The tissue movement can be caused by physiological processes internally in the organ such as the pulsation of the arteries. The tissue can also be externally excited by manually pushing the tissue or by using an electromechanical vibrating device. An alternative approach is to use the acoustic radiation force of an ultrasonic focused beam to generate displacements in the tissue with subsequent detection of the mechanical properties. One example of such an approach is the Acoustic Radiation Force Impulse (ARFI) method developed at Duke University [16]. In this technique, short duration acoustic pulses (push pulses) are used to generate small localized displacements deep in the tissue. These displacements are tracked by ultrasonic cross corre‐ lation, in a similar fashion as for the quasi-static elasticity imaging. The method has been

Another example is the innovative Supersonic Shear Imaging (SSI) method developed by the research group at the Laboratoire Ondes et Acoustique [20]. In SSI the acoustic radiation force is used to generate low-frequency shear waves (50-500 Hz) remotely in the tissue. The shear modulus of the tissue can be quantified by imaging the share wave propagation in the tissue by using ultrasound frame rates of several kHz. The method has been explored for diagnosis

For a more detailed overview about methods for ultrasound elasticity imaging and its clinical use we recommend to read the review papers by Wells and Liang [24] and Parker, Doyley and

investigated for imaging of focal liver lesions, prostate and breast [17-19].

time images.

32 Advancements and Breakthroughs in Ultrasound Imaging

as an option on commercial ultrasound systems.

of liver fibrosis, breast lesions and cornea [21-23].

Rubens [25].

In 1980, Carstensen and Muir published two papers describing the importance of nonlinear acoustics within the field of medical ultrasound imaging [26, 27]. These papers predicted and demonstrated nonlinear acoustical effects relevant for intensities and frequencies common in biomedical imaging. There has been an increasing interest with respect to nonlinear biomedical acoustics during the last 30 years. This interest was further escalated by the introduction of ultrasound contrast agents in the form of microbubbles and the study of these microbubbles was the main impetus for the introduction of the tissue harmonic imaging technique.

Nonlinear effects can be important in the forward wave propagation. The back-scattered pressure levels of the echoes are typically too low to induce any significant nonlinear effects. One source of nonlinear terms is produced by the deformation of tissue volume elements during compression and expansion with strongly curved phase fronts. It is, however, common to use transmit beams with relatively smooth phase fronts. Consequently, this nonlinear source is usually not the most dominant. The other important nonlinear source is nonlinear terms in the tissue elasticity and hence in the relation between acoustic pressure and tissue compres‐ sion/expansion. Nonlinear terms in the tissue elasticity are responsible for the fact that the tissue becomes stiffer during compression and softer during expansion. The compression also increases the mass density of the tissue, but this effect is inferior to the increased stiffness and the propagation velocity and will therefore be pressure dependent and will increase with increasing compressions and thus with increasing pressure. The resulting distortion of the transmit pressure field produces harmonic components which today are utilized in tissue harmonic imaging, especially in transcutaneous cardiac and abdominal imaging to suppress multiple scattering [28-31].

Ultrasound imaging is based on several assumptions, and one important assumption is that multiple scattering is neglected. For many organs, this approximation is valid. However, for the body wall, where larger variations in material parameters often are found, this assumption can be inadequate. Interfaces between soft tissue components with significant differences in material parameters give so strong echoes from the transmitted acoustic pulses that multiple scattering can get significant amplitudes. Such multiple scatterings are usually termed pulse reverberations [32, 33]. These reverberations reduce the ratio of the strongest to the weakest scatterer that can be detected in the neighborhood of each other, defined as the contrast resolution in the image. Reduced contrast resolution is in particular a problem when imaging hypo-echoic structures such as the heart chambers, the lumen of large blood vessels, some atherosclerotic lesions, cysts, some tumors, the gallbladder as well as in fetal imaging. The contact interface between the ultrasound transducer itself and the soft tissue is also a strong reflector enhancing the problem with multiple scattering.

Ultrasound contrast agents are made as a suspension of gas microbubbles encapsulated in thin stabilizing shells made from lipid or albumin. Typical bubble size is in the 1-5 μm range and the contrast bubbles are intravenously injected to increase the scattering from blood, which is weak compared to the scattering from soft tissues. Commercially available contrast bubbles are stable and small enough to enable transpulmonary passage and the blood half-life is typically in the range of 1-10 minutes. Scattering from microbubbles occurring within a liquid is resonant through an interaction between a co-oscillating liquid mass around the bubble and the bubble compression elasticity [34] with typical resonance frequencies of 1-7 MHz. With adequately flexible shells, the gas bubble has a very high compliance relative to the surround‐ ing blood and when driven by ultrasound pulses at frequencies below or around the bubble resonance frequency, large bubble radius excursions on the order of one micrometer is achieved due to mainly shear deformation and limited volume compression of the blood surrounding the bubble. This bubble radius displacement is then between one and two orders of magnitude larger than typical particle displacements obtained within soft tissues. The radius oscillation of a bubble may be obtained from the Rayleigh-Plesset equation [35, 36]:

$$\begin{array}{llll} \, \rho \left( a\ddot{a} + \frac{3}{2} \dot{a}^2 \right) = & \mathbf{-B} \langle a, \dot{a} \rangle \text{ - } p\_i(t) \end{array} \tag{1}$$

components are present in the scattered pressure from the bubble. This response is obtained with an incident drive pulse having a mechanical index equal to 0.07, which is very low compared to what is used for regular tissue imaging. At such low transmit pressure levels, the forward wave propagation will be close to linear and distortion of the transmit field due to nonlinear tissue elasticity will thus be very low. The harmonic components can then be used to differentiate bubble echoes from tissue echoes through Pulse Inversion and Amplitude Modulation pulsing schemes. In most clinical applications of ultrasound contrast agents, it is desirable to assess the micro-circulation or the tissue perfusion which cannot be done without the use of contrast agents and which often is related to various diseases. It is then necessary to obtain a strong suppression of the tissue signal for detection of the contrast bubble signal.

**Figure 5.** Numerical simulation of oscillation for a bubble with equilibrium radius of 2 μm and resonance frequency of 2.5 MHz. The upper panels show the drive pulse, the middle panels show the resulting bubble radius oscillation and the lower panels show the far-field component of the scattered pressure from the bubble. The left panels display the

scattering.

ߩ ቀܽܽሷ <sup>ଷ</sup> ଶ ܽሶ

be written

panel.

ܤሺܽǡ ܽሶሻ ൌ െቀ ଶఙ

బ ቁ ቀబ ቁ ଷ

Pulse Inversion methods [42-44].

0 0.1 (MPa)

> 1.6 1.8 2 2.2 2.4 (m)

> > �0.5 0 0.5 1

<sup>2</sup> <sup>3</sup> <sup>4</sup> <sup>5</sup> <sup>6</sup> �0.1

2 3 4 5 6

2 3 4 5 6

pulses in the time domain whereas the modulus of the Fourier Transform is displayed in the right panels.

ଶቁ ൌ െܤሺܽǡ ܽሶሻ െሺݐሻ (1)

between the ultrasound transducer itself and the soft tissue is also a strong reflector enhancing the problem with multiple

Ultrasound contrast agents are made as a suspension of gas microbubbles encapsulated in thin stabilizing shells made from lipid or albumin. Typical bubble size is in the 1-5 μm range and the contrast bubbles are intravenously injected to increase the scattering from blood, which is weak compared to the scattering from soft tissues. Commercially available contrast bubbles are stable and small enough to enable transpulmonary passage and the blood half-life is typically in the range of 1-10 minutes. Scattering from microbubbles occurring within a liquid is resonant through an interaction between a co-oscillating liquid mass around the bubble and the bubble compression elasticity [34] with typical resonance frequencies of 1-7 MHz. With adequately flexible shells, the gas bubble has a very high compliance relative to the surrounding blood and when driven by ultrasound pulses at frequencies below or around the bubble resonance frequency, large bubble radius excursions on the order of one micrometer is achieved due to mainly shear deformation and limited volume compression of the blood surrounding the bubble. This bubble radius displacement is then between one and two orders of magnitude larger than typical particle displacements obtained within soft tissues. The

where ߩ is the mass density of the surrounding liquid, ܽ is the bubble radius (where the time dependence has been omitted for convenience), ܤ is the pressure produced by the gas and the encapsulating shell, is the incident drive pressure and the dots represent differentiation with respect to time so that ܽሶ and ܽሷ represent the velocity and acceleration of the bubble wall, respectively. The terms on the left-hand side represent acceleration forces of the co-oscillation liquid mass whereas the terms on the right-hand side represent pressure terms due to gas and shell elasticity in addition to the drive pressure. The bubble pressure ܤ can

where the first term is the gas pressure and where ߢ is the polytropic exponent of the gas and ܽ is the equilibrium bubble radius. The second term ܵ is the pressure contribution from the encapsulating shell and is the ambient hydrostatic pressure. The fourth term accounts for surface tension due to the gas-liquid interface and the last term accounts for damping effects. When a contrast bubble is insonified by frequencies below or around its resonance frequency, the local nonlinear scattering from the contrast bubble is usually much larger than from soft tissues [37, 38]. This has resulted in several nonlinear ultrasound contrast agent detection techniques with the purpose to suppress the linear part of a received signal while maintaining as much as possible of the nonlinear part of a received signal. This is then used for low transmit pressure levels. The forward wave propagation is close to linear whereas the scattering from microbubbles can be highly nonlinear. Common techniques in use today are Pulse Inversion methods that detect even harmonic components [39, 40]. Amplitude Modulation methods are also in use [41], often in combination with

Ultrasound-Based Guidance and Therapy http://dx.doi.org/10.5772/55884 35

Figure 5. Numerical simulation of oscillation for a bubble with equilibrium radius of 2 m and resonance frequency of 2.5 MHz. Upper panel shows the drive pulse, middle panel shows the resulting bubble radius oscillation and lower panel shows the far-field component of the scattered pressure from the bubble. Left panel displays the pulses in the time domain whereas the modulus of the Fourier Transform is displayed in the right

The equations describing the bubble oscillations can be solved numerically. An example of a bubble with equilibrium radius of 2 m is shown in figure 5. An incident drive pulse with center frequency around 2 MHz is displayed in the time and frequency domain in the upper panel. In the middle panel, the resulting bubble radius oscillation is depicted and in the lower panel, the resulting normalized far-field component of the scattered pressure from the bubble is displayed. It can bee seen that the response is highly nonlinear and several harmonic components are present in the scattered pressure from the bubble. This response is obtained with an incident drive pulse having a mechanical index equal to 0.07, which is very low compared to what is used for regular tissue imaging. At such low transmit pressure levels, the forward wave propagation will be close to linear and distortion of the transmit field due to nonlinear tissue elasticity will thus be very low. The harmonic components can then be used to differentiate bubble

ሶ ߤ 

 (2)

radius oscillation of a bubble may be obtained from the Rayleigh-Plesset equation [35, 36]:

െ ܵሺܽǡ ܽሻ ଶఙ

<sup>0</sup> <sup>2</sup> <sup>4</sup> <sup>6</sup> <sup>8</sup> <sup>10</sup> �<sup>40</sup>

<sup>0</sup> <sup>2</sup> <sup>4</sup> <sup>6</sup> <sup>8</sup> <sup>10</sup> �<sup>40</sup>

Frequency (MHz)

�20 0 (dB)

�20 0 (dB)

(dB)

Time (s) <sup>0</sup> <sup>2</sup> <sup>4</sup> <sup>6</sup> <sup>8</sup> <sup>10</sup> �<sup>40</sup> �20 0

An example of the use of ultrasound contrast agents in relation to minimally invasive inter‐ ventions is radiofrequency ablation of liver tumors where contrast-enhanced ultrasound is used for improved detection and imaging of the lesions, for planning and guidance of multiple

SURF (Second order UltRasound Field) imaging is a nonlinear ultrasound imaging technique being developed in Trondheim [46-50]. It is based on transmission of dual frequency band pulse complexes consisting of a low frequency manipulation pulse and a high frequency imaging pulse that are co-propagating. Two transmit pulse complexes that may be used with the SURF technique are displayed in figure 6. With the use of conventional single frequency band transmit pulses, nonlinear effects are mainly restricted to the generation of harmonic components of the imaging pulse. With dual frequency band transmit pulses, other nonlinear effects also come into play. SURF imaging aims at further utilizing nonlinear acoustics for

For imaging of ultrasound microbubbles, conventional techniques relies on driving the bubble into strong nonlinear oscillations with the imaging pulse at relatively low mechanical indexes. This is typically feasible when the imaging frequency is below or around the bubble resonance frequency (as in the example of figure 5) and conventional contrast agents typically have resonance frequencies below 7 MHz. However, when the imaging frequency is above the bubble resonance frequency a much higher mechanical index is required to obtain significant

needle electrodes and finally for immediate evaluation of the treatment [45].

improved imaging of various tissues and ultrasound contrast agents.

where *ρ* is the mass density of the surrounding liquid, *a* is the bubble radius (where the time dependence has been omitted for convenience), *B* is the pressure produced by the gas and the encapsulating shell, *pi* is the incident drive pressure and the dots represent differentiation with respect to time so that *a*˙ and *a* ¨ represent the velocity and acceleration of the bubble wall, respectively. The terms on the left-hand side represent acceleration forces of the co-oscillation liquid mass whereas the terms on the right-hand side represent pressure terms due to gas and shell elasticity in addition to the drive pressure. The bubble pressure *B* can be written

$$B\begin{pmatrix} a \ \dot{a} \end{pmatrix} = -\left(p\_0 + \frac{2\sigma}{a\_0}\right)\left(\frac{a\_0}{a}\right)^{3\times} \ \ \ \ \ \mathcal{S}\begin{pmatrix} a\_0 \ a \end{pmatrix} + \ \ p\_0 + \frac{2\sigma}{a} + \ \mu \frac{a}{a} \tag{2}$$

where the first term is the gas pressure and where *κ* is the polytropic exponent of the gas and *a*0 is the equilibrium bubble radius. The second term *S* is the pressure contribution from the encapsulating shell and *p*<sup>0</sup> is the ambient hydrostatic pressure. The fourth term accounts for surface tension due to the gas-liquid interface and the last term accounts for damping effects. When a contrast bubble is insonified by frequencies below or around its resonance frequency, the local nonlinear scattering from the contrast bubble is usually much larger than from soft tissues [37, 38]. This has resulted in several nonlinear ultrasound contrast agent detection techniques with the purpose to suppress the linear part of a received signal while maintaining as much as possible of the nonlinear part of a received signal. This is then used for low transmit pressure levels. The forward wave propagation is close to linear whereas the scattering from microbubbles can be highly nonlinear. Common techniques in use today are Pulse Inversion methods that detect even harmonic components [39, 40]. Amplitude Modulation methods are also in use [41], often in combination with Pulse Inversion methods [42-44].

The equations describing the bubble oscillations can be solved numerically. An example of a bubble with equilibrium radius of 2 μm is shown in figure 5. An incident drive pulse with center frequency around 2 MHz is displayed in the time and frequency domain in the upper panel. In the middle panel, the resulting bubble radius oscillation is depicted and in the lower panel, the resulting normalized far-field component of the scattered pressure from the bubble is displayed. It can bee seen that the response is highly nonlinear and several harmonic between the ultrasound transducer itself and the soft tissue is also a strong reflector enhancing the problem with multiple

Ultrasound contrast agents are made as a suspension of gas microbubbles encapsulated in thin stabilizing shells made from lipid or albumin. Typical bubble size is in the 1-5 μm range and the contrast bubbles are intravenously injected to increase the scattering from blood, which is weak compared to the scattering from soft tissues. Commercially available contrast bubbles are stable and small enough to enable transpulmonary passage and the blood half-life is typically in the range of 1-10 minutes. Scattering from microbubbles occurring within a liquid is resonant through an interaction between a co-oscillating liquid mass around the bubble and the bubble compression elasticity [34] with typical resonance frequencies of 1-7 MHz. With adequately flexible shells, the gas bubble has a very high compliance relative to the surrounding blood and when driven by ultrasound pulses at frequencies below or around the bubble resonance frequency, large bubble radius excursions on the order of one micrometer is achieved due to mainly shear deformation and limited volume compression of the blood surrounding the bubble. This bubble radius displacement is then between one and two orders of magnitude larger than typical particle displacements obtained within soft tissues. The

where ߩ is the mass density of the surrounding liquid, ܽ is the bubble radius (where the time dependence has been omitted for convenience), ܤ is the pressure produced by the gas and the encapsulating shell, is the incident drive pressure and the dots represent differentiation with respect to time so that ܽሶ and ܽሷ represent the velocity and acceleration of the bubble wall, respectively. The terms on the left-hand side represent acceleration forces of the co-oscillation liquid mass whereas the terms on the right-hand side represent pressure terms due to gas and shell elasticity in addition to the drive pressure. The bubble pressure ܤ can

where the first term is the gas pressure and where ߢ is the polytropic exponent of the gas and ܽ is the equilibrium bubble radius. The second term ܵ is the pressure contribution from the encapsulating shell and is the ambient hydrostatic pressure. The fourth term accounts for surface tension due to the gas-liquid interface and the last term accounts for damping effects. When a contrast bubble is insonified by frequencies below or around its resonance frequency, the local nonlinear scattering from the contrast bubble is usually much larger than from soft tissues [37, 38]. This has resulted in several nonlinear ultrasound contrast agent detection techniques with the purpose to suppress the linear part of a received signal while maintaining as much as possible of the nonlinear part of a received signal. This is then used for low transmit pressure levels. The forward wave propagation is close to linear

ሶ ߤ 

 (2)

radius oscillation of a bubble may be obtained from the Rayleigh-Plesset equation [35, 36]:

െ ܵሺܽǡ ܽሻ ଶఙ

scattering.

ߩ ቀܽܽሷ <sup>ଷ</sup> ଶ ܽሶ

be written

ܤሺܽǡ ܽሶሻ ൌ െቀ ଶఙ

బ ቁ ቀబ ቁ ଷ

Pulse Inversion methods [42-44].

ଶቁ ൌ െܤሺܽǡ ܽሶሻ െሺݐሻ (1)

is resonant through an interaction between a co-oscillating liquid mass around the bubble and the bubble compression elasticity [34] with typical resonance frequencies of 1-7 MHz. With adequately flexible shells, the gas bubble has a very high compliance relative to the surround‐ ing blood and when driven by ultrasound pulses at frequencies below or around the bubble resonance frequency, large bubble radius excursions on the order of one micrometer is achieved due to mainly shear deformation and limited volume compression of the blood surrounding the bubble. This bubble radius displacement is then between one and two orders of magnitude larger than typical particle displacements obtained within soft tissues. The radius

oscillation of a bubble may be obtained from the Rayleigh-Plesset equation [35, 36]:

<sup>2</sup> *<sup>a</sup>*˙2) <sup>=</sup> - *<sup>B</sup>*(*a*, *<sup>a</sup>*˙) - *pi*

where *ρ* is the mass density of the surrounding liquid, *a* is the bubble radius (where the time dependence has been omitted for convenience), *B* is the pressure produced by the gas and the

respectively. The terms on the left-hand side represent acceleration forces of the co-oscillation liquid mass whereas the terms on the right-hand side represent pressure terms due to gas and shell elasticity in addition to the drive pressure. The bubble pressure *B* can be written

where the first term is the gas pressure and where *κ* is the polytropic exponent of the gas and *a*0 is the equilibrium bubble radius. The second term *S* is the pressure contribution from the encapsulating shell and *p*<sup>0</sup> is the ambient hydrostatic pressure. The fourth term accounts for surface tension due to the gas-liquid interface and the last term accounts for damping effects. When a contrast bubble is insonified by frequencies below or around its resonance frequency, the local nonlinear scattering from the contrast bubble is usually much larger than from soft tissues [37, 38]. This has resulted in several nonlinear ultrasound contrast agent detection techniques with the purpose to suppress the linear part of a received signal while maintaining as much as possible of the nonlinear part of a received signal. This is then used for low transmit pressure levels. The forward wave propagation is close to linear whereas the scattering from microbubbles can be highly nonlinear. Common techniques in use today are Pulse Inversion methods that detect even harmonic components [39, 40]. Amplitude Modulation methods are

The equations describing the bubble oscillations can be solved numerically. An example of a bubble with equilibrium radius of 2 μm is shown in figure 5. An incident drive pulse with center frequency around 2 MHz is displayed in the time and frequency domain in the upper panel. In the middle panel, the resulting bubble radius oscillation is depicted and in the lower panel, the resulting normalized far-field component of the scattered pressure from the bubble is displayed. It can bee seen that the response is highly nonlinear and several harmonic


is the incident drive pressure and the dots represent differentiation with

¨ represent the velocity and acceleration of the bubble wall,

2*σ <sup>a</sup>* <sup>+</sup> *<sup>μ</sup> <sup>a</sup>*˙

(*t*) (1)

*<sup>a</sup>* (2)

*ρ*(*aa* ¨ + 3

*B*(*a*, *a*˙)= - ( *p*<sup>0</sup> +

2*σ a*0 )( *a*0 *a* )3*κ*

also in use [41], often in combination with Pulse Inversion methods [42-44].

encapsulating shell, *pi*

respect to time so that *a*˙ and *a*

34 Advancements and Breakthroughs in Ultrasound Imaging

Figure 5. Numerical simulation of oscillation for a bubble with equilibrium radius of 2 m and resonance frequency of 2.5 MHz. Upper panel shows the drive pulse, middle panel shows the resulting bubble radius oscillation and lower panel shows the far-field component of the scattered pressure from the bubble. Left panel displays the pulses in the time domain whereas the modulus of the Fourier Transform is displayed in the right panel. **Figure 5.** Numerical simulation of oscillation for a bubble with equilibrium radius of 2 μm and resonance frequency of 2.5 MHz. The upper panels show the drive pulse, the middle panels show the resulting bubble radius oscillation and the lower panels show the far-field component of the scattered pressure from the bubble. The left panels display the pulses in the time domain whereas the modulus of the Fourier Transform is displayed in the right panels.

components are present in the scattered pressure from the bubble. This response is obtained with an incident drive pulse having a mechanical index equal to 0.07, which is very low compared to what is used for regular tissue imaging. At such low transmit pressure levels, the forward wave propagation will be close to linear and distortion of the transmit field due to nonlinear tissue elasticity will thus be very low. The harmonic components can then be used to differentiate bubble echoes from tissue echoes through Pulse Inversion and Amplitude Modulation pulsing schemes. In most clinical applications of ultrasound contrast agents, it is desirable to assess the micro-circulation or the tissue perfusion which cannot be done without the use of contrast agents and which often is related to various diseases. It is then necessary to obtain a strong suppression of the tissue signal for detection of the contrast bubble signal. The equations describing the bubble oscillations can be solved numerically. An example of a bubble with equilibrium radius of 2 m is shown in figure 5. An incident drive pulse with center frequency around 2 MHz is displayed in the time and frequency domain in the upper panel. In the middle panel, the resulting bubble radius oscillation is depicted and in the lower panel, the resulting normalized far-field component of the scattered pressure from the bubble is displayed. It can bee seen that the response is highly nonlinear and several harmonic components are present in the scattered pressure from the bubble. This response is obtained with an incident drive pulse having a mechanical index equal to 0.07, which is very low compared to what is used for regular tissue imaging. At such low transmit pressure levels, the forward wave propagation will be close to linear and distortion of the transmit field due to nonlinear tissue elasticity will thus be very low. The harmonic components can then be used to differentiate bubble

An example of the use of ultrasound contrast agents in relation to minimally invasive inter‐ ventions is radiofrequency ablation of liver tumors where contrast-enhanced ultrasound is used for improved detection and imaging of the lesions, for planning and guidance of multiple needle electrodes and finally for immediate evaluation of the treatment [45].

SURF (Second order UltRasound Field) imaging is a nonlinear ultrasound imaging technique being developed in Trondheim [46-50]. It is based on transmission of dual frequency band pulse complexes consisting of a low frequency manipulation pulse and a high frequency imaging pulse that are co-propagating. Two transmit pulse complexes that may be used with the SURF technique are displayed in figure 6. With the use of conventional single frequency band transmit pulses, nonlinear effects are mainly restricted to the generation of harmonic components of the imaging pulse. With dual frequency band transmit pulses, other nonlinear effects also come into play. SURF imaging aims at further utilizing nonlinear acoustics for improved imaging of various tissues and ultrasound contrast agents.

For imaging of ultrasound microbubbles, conventional techniques relies on driving the bubble into strong nonlinear oscillations with the imaging pulse at relatively low mechanical indexes. This is typically feasible when the imaging frequency is below or around the bubble resonance frequency (as in the example of figure 5) and conventional contrast agents typically have resonance frequencies below 7 MHz. However, when the imaging frequency is above the bubble resonance frequency a much higher mechanical index is required to obtain significant

the tracking data could be directed directly into the scanner and the ultrasound frames could be used off-line, e.g. to generate a 3D volume from the tagged 2D frames). The traditional way of getting real-time access to ultrasound frames is to connect the analog output (e.g., composite video, S-video) of the ultrasound scanner to a frame-grabbing card on the navigation computer. Using the analog output might affect the image quality due to the double digital-to-analog-todigital conversion and no metadata (e.g. depth) follow the ultrasound images. Alternatively digital data can be streamed directly from the ultrasound scanner and into the navigation computer. Traditionally this has required some kind of research collaboration between the ultrasound manufacturer and the user but open ultrasound scanners are becoming available (e.g. the Ultrasonix scanner). These systems usually provide just a one-way streaming interface but two-way communication protocols where the scanner can be controlled (e.g. depth) by the navigation system exists making more integrated solutions possible (figure 7). Either way, the protocol (or interface / API) used is typically proprietary, although proposals for real-time standards are starting to emerge (e.g. OpenIGTLink, DICOM in surgery (WG24)). When the link between the ultrasound scanner and navigation system is digital, ultrasound data at different stages in the processing chain on the scanner can be transferred (e.g. scan-converted, scan-line and RF-data). Furthermore, a digital streaming interface will be required in order to use the real-time 3D scanners that are now becoming available also for navigation. It's difficult to capture the 3D content in the scanner display using a frame grabber so the data needs to be transferred in real-time or tagged with a tracking reference on the ultrasound scanner.

Ultrasound-Based Guidance and Therapy http://dx.doi.org/10.5772/55884 37

**Figure 7.** Streaming ultrasound data into the navigation system. The interface can either be analog using a frame grabber or digital using a direct link and a proprietary protocol. A digital interface can either be one-way (i.e. stream‐ ing) or two-way (i.e. optionally control the scanner from the navigation system as well). In any case the image stream

In order to use ultrasound to guide surgical procedures the ultrasound probe must be tracked. Several tracking technologies have been proposed over the years (mechanical, acoustical,

must be tagged with tracking data and in order to do that the two streams need to be synchronized.

**3.2. Tracking of ultrasound probes**

**Figure 6.** Example of SURF transmit pulse complexes where a low frequency manipulation pulse at 1 MHz is co-propa‐ gating with a high frequency imaging pulse at 10 MHz. The high frequency imaging pulse is in the left and right panel placed at low and high manipulation pressure, respectively.

nonlinear back-scattering from the bubble. At higher mechanical indexes the tissue will also respond nonlinearly and it then becomes difficult to differentiate the tissue signal from the bubble signal. For contrast imaging at high frequencies, such as 10 – 30 MHz, that can be used in minimal invasive interventions where the probe can be close to the object being imaged, conventional contrast imaging techniques often have limitations. The dual band SURF technique then has some advantages where the low frequency manipulation pulse can be tuned to match the bubble resonance frequency (typically around 2-3 MHz) whereas the high frequency imaging pulse can be optimized for the object being imaged and can for example be 20 MHz. The low frequency then manipulates the bubble oscillation and back-scattering which is interrogated by the high frequency pulse. The high frequency imaging pulse is hence decoupled from the resonance properties of the contrast bubbles.
