Advances In Image Quality And Image Presentation

Research in the field of US has been always aimed at obtaining high spatial resolution, good penetration of the US beam, and reduction of signal-to-noise ratio. This has been possible from basic research in the behavior of US in tissues, the physical principles underlying piezoelectric materials, and from progress in the construction and miniaturization of electronic components. The new developments include modern piezoelectric materials with lower characteristic acoustic impedance and greater electromechanical coupling coefficients, production of multiple layer transducers of ceramic materials of variable thickness and shape, better understanding of the behavior of small transducers, and the development of a variety of pulse-characteristic modulation processes that allow accurate control of transmission frequency, amplitude, phase, and pulse length (1-4). These advances have increased the ability of US in demonstrating anatomic and pathological details (Fig. 1). In the following section,

Figure 1 Gray scale US of renal allografts using a broadband, high-frequency linear transducer. (A) Normal appearance of a renal pyramid. The excellent spatial and contrast resolution allows differentiation between the outer (arrowheads) and inner (*) medulla. Open arrow indicates an arcuate vessel. (B) Renal allograft infection. High-resolution US shows thickening of the wall of a renal calix (curved arrows). Abbreviation: US, ultrasonography.

Figure 1 Gray scale US of renal allografts using a broadband, high-frequency linear transducer. (A) Normal appearance of a renal pyramid. The excellent spatial and contrast resolution allows differentiation between the outer (arrowheads) and inner (*) medulla. Open arrow indicates an arcuate vessel. (B) Renal allograft infection. High-resolution US shows thickening of the wall of a renal calix (curved arrows). Abbreviation: US, ultrasonography.

new techniques that caused marked improvement in the quality of US imaging will be discussed. These include tissue harmonic imaging, compound imaging, the extended field-of-view technique, three-dimensional (3-D) imaging, and the use of US contrast media.

Tissue Harmonic Imaging

Increased knowledge in acoustic harmonics has led to a marked improvement in general ultrasonic imaging. The use of tissue harmonic imaging has led to a marked improvement in image quality, particularly in large, difficult-to-scan patients (1,3,4).

At the level of the kidneys and in patients with pelvic masses, harmonic imaging facilitates the differentiation between solid and cystic lesions (Fig. 2), allows better delineation of intracystic septa and vegetations (Fig. 3), and improves visualization of stones within dilated ureters by providing better delineation of the acoustic shadow (Fig. 4). At present, tissue harmonic imaging is suggested as the standard setting to be used in abdominal and pelvic US examination, especially when a fluid-filled lesion is under evaluation and its characteristics have to be analyzed (5-8).

Compound Imaging

Compound imaging is a relatively new technique that allows combining echoes obtained from US beams oriented along different directions. Electronic steering of US beams from an array transducer is used to image the same tissue multiple times from different directions; then the echoes from these multiple acquisitions are averaged together into a single composite image (1,4). This is possible at a real-time rate, although the frame-rate is slightly lower than that of conventional imaging. Compound imaging improves image quality by reducing speckle, without any loss in spatial resolution, and can provide better delineation of lesion contours (Fig. 5). Another approach to compounding, called ''frequency compound,'' uses simultaneous emission of two US

Figure 2 Comparison between (A) conventional gray scale US and (B) tissue harmonic imaging. A small renal cyst (arrow) is better characterized using tissue harmonic imaging due to reduced image artifacts. Abbreviation: US, ultrasonography.

beams of different frequencies, again resulting in better image quality, with less noise, reduced speckle, and good contrast resolution (7,8).

In clinical practice, the simultaneous use of tissue harmonic imaging and image compounding is often regarded as the standard setting for abdominal and pelvic US examinations. It must be noted, however, that the acoustic shadow posterior to small calcifications or stones can be less evident when using the image compounding technique, and this can be a drawback in some cases. Hence, switching between standard setting and compound imaging to allow delineation of acoustic shadows is suggested in difficult cases.

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Figure 3 Comparison between (A) conventional gray scale US and (B) tissue harmonic imaging. The internal septa of this small complex renal cyst are better visible using tissue harmonic imaging because of reduced image artifacts leading to increased contrast resolution between the liquid and solid components of the lesion. Abbreviation: US, ultrasonography.

Figure 3 Comparison between (A) conventional gray scale US and (B) tissue harmonic imaging. The internal septa of this small complex renal cyst are better visible using tissue harmonic imaging because of reduced image artifacts leading to increased contrast resolution between the liquid and solid components of the lesion. Abbreviation: US, ultrasonography.

Figure 4 Comparison between (A) conventional gray scale US and (B) tissue harmonic imaging. Reduced artifacts allow better identification of an uretheral stone (arrow) and of its acoustic shadow (arrowheads) in the harmonic image. Abbreviation: US, ultrasonography.

Extended Field-of-View

US images have a field-of-view that is limited by the probe width. During the study, the examiner moves the probe over the area of interest to acquire information on large volumes of tissues and reconstructs in his/her mind the spatial relationships between parts of anatomy by memorizing many small image frames (9). This is a distinct disadvantage of US compared to other imaging methods and is a major drawback in conveying the information of the study to clinicians. The extended field-of-view technique has been developed to overcome these limitations and allows the reconstruction of wide images by progressive addition of data during a hand sweep with the conventional real-time small probe (Fig. 6). The information about position is obtained directly from the ultrasound images themselves, using an image registration-based position-sensing technique, estimating probe position by combining multiple local motion vectors (1,9).

A number of papers have shown the clinical usefulness of this technique in the evaluation of large pelvic and abdominal lesions as well as in providing anatomic

Figure 5 Comparison between (A) conventional gray scale US and (B) compound imaging. Improved image quality of compound imaging allows better visualization of a small renal tumor (*). A thin peripheral hypoechoic rim (arrowheads), which is visible only in the compound image, suggests presence of tumor pseudocapsule. Abbreviation: US, ultrasonography.

Figure 5 Comparison between (A) conventional gray scale US and (B) compound imaging. Improved image quality of compound imaging allows better visualization of a small renal tumor (*). A thin peripheral hypoechoic rim (arrowheads), which is visible only in the compound image, suggests presence of tumor pseudocapsule. Abbreviation: US, ultrasonography.

Figure 6 Extended field-of-view images of the scrotal content in two patients. (A) Simultaneous visualization of the testis (T) and of a large cyst of the head of the epididymis (*). (B) Simultaneous visualization of right (R) and left (L) testes and of hydrocele surrounding the left testis.

context that was not possible with the small field-of-view of the real-time probe (10,11). This is a distinct advantage when explaining the ultrasound images to referring clinicians. It must be noted that the accuracy of measurements obtained with the extended field-of-view system has been proven accurate and repeatable (9,12), thus providing the capability to follow-up volume changes of large lesions during therapy.

3-D Imaging

The development of high-speed computing and the increase in storage capacity of hardware have opened the possibility of applying 3-D technologies to diagnostic ultrasound. Presenting the entire volume of data in a single image can overcome the problems in understanding difficult anatomy, clarify exact spatial relationships, and help sharing information with referring clinicians and patients (1,13).

To acquire volume data sets, it is necessary to obtain a series of contiguous 2-D image planes of the volume under evaluation. The position of each slice within the volume has to be precisely assessed, and this can be obtained through different techniques (13).

Special transducers have been developed integrating the position sensing system within the transducer housing. These ''volume transducers'' are larger than standard probes and can be more difficult to use; however, they can provide exact knowledge of each scan plane, eliminating distortions in the resultant images. After each slice the transducer plane is automatically moved to the next location, which is then exactly known by the machine. These transducers are usually integrated with the ultrasound equipment, and 3-D images can be immediately displayed after acquisition.

Position sensor devices can be attached to 2-D conventional transducers to obtain data about scan position. At present, small electromagnetic sensors mounted on the transducer are the most used technique for this purpose and, although susceptible to distortion from adjacent metallic objects, have proven able in producing accurate measurements.

The display of 3-D information gathered by US techniques is not simple, and it may be difficult to separate adjacent structures. 3-D ultrasound images, in fact, cannot be classified and processed like computed tomography (CT) and magnetic resonance imaging (MRI) data because they do not represent well-defined parameters, such as density, but are rather a measure of how the acoustic impedance changes as sound waves travel through tissues.

Slice projection from the volume data of images of arbitrary orientation is the most common method to review 3-D ultrasound data. This technique allows retrospective evaluation of anatomy along planes different from the one along which the data themselves have been obtained and has been proven to be useful in a variety of clinical settings.

There are many clinical applications of 3-D ultrasound imaging, particularly in assessing intrauterine fetal anatomy and gynecological imaging (13). In the urinary tract, volume estimation of the urinary bladder with 3-D ultrasound has been shown to be accurate, reliable, and clinically useful (14-16). Volume measurements have also proven accurate in assessing prostatic cancer and benign hyperplasia, with better identification of extraglandular spread (17,18). However, better identification did not result in increased accuracy in cancer staging (17). Ultrasound 3-D imaging can also be used for accurate measurements of renal volume and for monitoring progress in patients with renal diseases. In addition, it can offer useful anatomical information about the renal pelvis and its relation to the branches of the renal artery (19-21).

Ultrasound Contrast Media

The possibility of enhancing visualization of the vascular system with US reflectors after the injection of a variety of fluids was originally described in 1968 (22). Shortly afterwards, it was understood that the source of the observed intravascular echoes was microbubbles developing during the injection process. Since then, the pharmaceutical industry has worked to develop safe and nontoxic, intravenously injectable products made of microbubbles stable enough to cross the pulmonary capillary bed after injection in a peripheral vein and to provide vascular enhancement for the whole duration of the clinical examination (23). The technology adopted has been that of encapsulated bubbles of gas, smaller than the red blood cells (24), with the capability to flow freely in the vascular system. A variety of gases have been used, from air to less diffusible compounds, such as perfluorocarbons or sulfur hexafluor-ide. A variety of shells with different thicknesses and stiffnesses have been used to encapsulate the microbubbles.

US contrast agents behave as an active source of sound, modifying the characteristic signature of the echo from blood. When properly insonated with a highpower ultrasound beam, microbubbles collapse, producing a high intensity, broadband transient signal. When the power of the ultrasound beam is lower, microbubbles undergo complex oscillation in the ultrasound field and work by resonating, rapidly contracting and expanding in response to the pressure changes of the sound wave, producing multiple harmonic signals.

Microbubble contrast agents are neither filtered by the kidney nor able to enter the interstitial spaces and act as "blood pool agents'' until metabolized. However, some have recently been shown to exhibit specific hepatosplenic uptake after their disappearance from the blood pool (24-26).

US contrast agents were originally developed to increase vascular signals during Doppler studies in "difficult" patients and rescue otherwise failed examinations. As a consequence, early clinical experience in uroradiology was mainly focused on visualization of renal arteries and identifying small or deep vessels within parench-ymal lesions. These early clinical studies have demonstrated that the success rate of Doppler visualization of the main renal arteries and of accessory renal arteries improves significantly after the administration of microbubbles (27-30).

A variety of softwares have been developed by the manufacturers of US equipment to detect signals from microbubbles. There are basically two imaging approaches that can be used: "destructive contrast-specific modes,'' which collect the signal from bubble destruction produced using a high-power US beam, and "nondestructive contrast-specific modes,'' which collect the harmonic signals from bubble insonation with US beams of lower power.

Contrast-specific destructive modes allow excellent depiction of renal vascular-ity and perfusion defects from different causes, such as renal infection, focal ischemia, or trauma, and differential diagnosis between renal neoplasm and pseudotumors (31-33). Heterogeneous contrast enhancement usually suggests malignancy. In addition, cystic renal tumors can be differentiated from complex benign cystic masses in most cases when contrast enhancement is appreciable within the lesion wall (Fig. 7) or within intracystic septa (33). However, inflammatory cysts can occasionally simulate malignancy (34).

Contrast-specific destructive modes have been also used successfully to evaluate cancer of the prostate gland. Detection of isoechoic tumors and cancers of both the peripheral and central regions of prostate improves using contrast-specific modes, with better guidance to biopsy (35-38).

A technique that allows evaluation of vesico-ureteral reflux has also been developed. After catheterization, the bladder is filled in with saline until the patient has the urge to micturate, and then the US contrast medium is added. Reflux is diagnosed when microbubbles are detected in the ureter or the renal pelvis. The results show good correlation with conventional micturating cystourethrogram and with radionuclide studies (39-42).

The major disadvantage of contrast-specific destructive modes is that intermittent scanning with a limited number of insonations is needed to minimize bubble destruction. Hence, prolonged evaluation of contrast enhancement cannot be performed. This limitation can be overcome by performing nondestructive low-acoustic

Figure 7 Small cystic renal tumor (2 cm). (A) Conventional power Doppler evaluation shows no vascular signals in the tumor wall. (B) Power Doppler image obtained with the same imaging parameters during Levovist infusion shows vascularity of tumor wall.

power US scanning after intravenous administration of US contrast agents. This technique allows imaging both kidneys in real time with excellent evaluation of vas-cularity and reduces artifacts even at unfavorable Doppler angles. Clinical evidence is accumulating that nondestructive scanning is more effective than destructive methods to evaluate all renal pathologies (43,44). Real-time imaging improves the visualization of small focal renal lesions and of perfusion defects (Fig. 8). Improved temporal resolution may allow visualization of segmental areas with delayed enhancement, permitting the diagnosis of segmental renal artery stenosis (Fig. 9) or atheroembolic renal disease. Detection of isoechoic renal tumors remains very difficult because these lesions often appear hyperechoic only in the early arterial phase (Fig. 10), while in the other vascular phases they are isoechoic to the normal renal parenchyma (43,44). However, identification of such tumors is better with nondestructive modes when compared with the destructive imaging mode. Moreover, real-time contrast-specific imaging is effective in improving the sonographic visualization of tumoral pseudocapsule, which appears after microbubble injection as a rim of perilesional enhancement, increasing in the latter phase of the examination (45).

Microbubble contrast agents can also be used as vascular tracers to assess renal perfusion. Preliminary studies performed with Doppler techniques on renal allo-grafts showed that measuring the arteriovenous transit time after bolus injection of microbubbles allows differentiation between normally functioning grafts and kidneys with biopsy-proven rejection (46). Other preliminary investigations showed that analysis of time-intensity curves drawn from the renal cortex after a bolus injection of microbubbles can improve identification of hemodynamically significant renal artery stenosis (47,48).

Renal perfusion can also be assessed using a different approach. If micro-bubbles are administered with infusion technique, a steady state is reached in which contrast blood concentration can be considered to be constant. When this state has

Figure 8 Seventy-two-year-old diabetic patient who presented with left flank pain. (A) Color Doppler US (shown here in gray scale): a hypoperfused area at the upper pole of the left kidney is suspected (*), but with low diagnostic confidence due to unfavorable Doppler angle. Microbubble administration allows to confirm with high diagnostic confidence that the upper portion of the left kidney is ischemic (*). (B) Note excellent contrast enhancement with high spatial resolution of the remaining portions of the left kidney (arrowheads) and of the spleen (S). Abbreviation: US, ultrasonography.

Figure 8 Seventy-two-year-old diabetic patient who presented with left flank pain. (A) Color Doppler US (shown here in gray scale): a hypoperfused area at the upper pole of the left kidney is suspected (*), but with low diagnostic confidence due to unfavorable Doppler angle. Microbubble administration allows to confirm with high diagnostic confidence that the upper portion of the left kidney is ischemic (*). (B) Note excellent contrast enhancement with high spatial resolution of the remaining portions of the left kidney (arrowheads) and of the spleen (S). Abbreviation: US, ultrasonography.

Figure 9 Segmental renal artery stenosis of the right kidney. (A) Contrast enhanced CT shows delayed contrast enhancement in the affected portion of the kidney (*) compared with the normally perfused portion (arrowheads). (B and C) Contrast enhanced US. (B) Ten seconds after microbubble injection a similar delayed enhancement can be appreciated in the territory perfused by stenotic artery (*), compared with the normally perfused portion (arrowheads). (C) Forty seconds after microbubble administration renal vascularity appears normal. Abbreviations: CT, computed tomography; US, ultrasonography.

Figure 9 Segmental renal artery stenosis of the right kidney. (A) Contrast enhanced CT shows delayed contrast enhancement in the affected portion of the kidney (*) compared with the normally perfused portion (arrowheads). (B and C) Contrast enhanced US. (B) Ten seconds after microbubble injection a similar delayed enhancement can be appreciated in the territory perfused by stenotic artery (*), compared with the normally perfused portion (arrowheads). (C) Forty seconds after microbubble administration renal vascularity appears normal. Abbreviations: CT, computed tomography; US, ultrasonography.

been reached, the bubbles within an imaged slice can be destroyed by applying a high-intensity frame to create an inflow void or ''negative bolus'' (49,50). When the next destructive beam is transmitted, the intensity of the echoes depends on the number of bubbles that have flown into the slice and so increases with longer intervals. If the process is repeated at a series of intervals, a time-intensity curve can be created. The slope of this curve is related to the speed of blood moving into the slice, the maximum signal intensity level relates to the vascular volume, and their product reflects the volume flow rate and approximates perfusion (49,50). This method to evaluate perfusion can be applied to any tissue. Attempts have been made to use it to detect changes in renal blood flow in animals with flow-limiting renal artery stenosis and

Figure 10 Small renal tumor. Contrast enhanced US. (A) During the early arterial phase (18 seconds after microbubble injection) the lesion (curved arrows) appears hyperechoic. (B) Forty seconds after microbubble injection the lesion (curved arrows) appears nearly iso-echoic to kidney. A central nonenhancing area (arrowhead), probably necrotic, is appreciable in both vascular phases. Abbreviation: US, ultrasonography.

Figure 10 Small renal tumor. Contrast enhanced US. (A) During the early arterial phase (18 seconds after microbubble injection) the lesion (curved arrows) appears hyperechoic. (B) Forty seconds after microbubble injection the lesion (curved arrows) appears nearly iso-echoic to kidney. A central nonenhancing area (arrowhead), probably necrotic, is appreciable in both vascular phases. Abbreviation: US, ultrasonography.

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