This essay has been submitted by a student. This is not an example of the work written by our professional essay writers.
If 2 or more US source simultaneously send a wave across the same volume of tissues, it is the net inflow or outflow of molecules that determines the resultant sound pressures in that region this is called constructive interference. If they are in the opposite direction this is called destructive interference.
If a number of small transducers elements arranged in a line simultaneously produce small sound waves, each wave will grow the same distance from the probe in a given time; they will form a parallel ray forming constructive interference making a beam of sound.
The number of pressure peaks generated in one second is termed the frequency of the wave obtained from the reciprocal of time between equivalent points on the wave. Determination of the frequency of the pulse at all other similar points on the wave gives rise to a spectrum of frequencies.
Intensity & Attenuation
The movement of molecules in the centre of the beam is more vigorous in the edge with high power. The power measured over a small area is called the intensity of the beam at that point expressed in Watts per square meter. Attenuation describes loss of intensity of a pulse as it travels through the tissues. It occurs by absorption of rays, scattering which is reflection from small structures.
Each time an interface between different tissues is encountered by a pulse through the tissues, a small proportion of energy is reflected as an echo. Fat to muscle interface reflect 1% of sound energy. Fluid to tissue interface is highly reflective. Gas to tissue are 99.99% reflective. So it is essential to eliminate any air the probe & the patient.
The proportion of sound energy reflected at a boundary depends on the difference in the characteristic impedance of the tissue forming the interface. It is a measure of the velocity imparted to the molecules as they respond to a rise in sound pressure. It is determined by elasticity & density of tissues.
Specular Reflection, Scatter & Shadowing
Specular or mirror like reflection occurs when the reflector is a relatively extensive surface such as capsule of an organ or surface of the placenta, bladder or the fetus. Scatter is generated by parenchymal structures. Shadowing occurs when an echo is so large that the sound continuing into the tissues is highly attenuated, an acoustic shadow is casted.
Basics of 3 Dimensional (3D) Ultrasonography
The aim of three dimensional ultrasound is the representation of the morphology of an organ of interest (Merz et al., 1998).
To achieve these results the following results are necessary (Leveine et al., 1992):
Echo data processing along an ultrasound beam.
Movement of an ultrasound beam over the area of interest. The three data set is created by the translation/rotation of the axis from the ultrasound beam and the flight time of the reflected sound waves which is converted into distance information by the assumption of the speed of sound within the volume of interest.
Storage of data & gap filling procedure.
Data representation and visualization.
Echo data processing is well known from 2DUS.the reflected ultrasound waves can be processed according to their amplitude (B-mode imaging), their frequency shift compared to the frequency that has been transmitted (Doppler imaging), their energy in spectral components that are shifted by moving targets (angiography, power Doppler) or their amplitude on harmonic components of the reflected echo data (harmonic imaging). All modalities of 2DUS can also be applied to 3DUS (Kossof, 1995).
Acquisition of 3D Images
Acquisition of the information needed to produce a three dimensional ultrasound image depends on obtaining echo data throughout the volume of interest (Fenster & Downey, 1996).
To produce proper images, these factors must be considered & optimized. First the scanning technique must be fast or gated to avoid artifacts and distortions caused by involuntary, respiratory or cardiac motion. Second, the geometry of scanning technique must be known accurately or calibrated correctly to avoid geometric distortions and measurement errors. Third, the scanning technique must be easy to use so that it does not interfere with the patient exam or it is not convenient to use. Four different approaches are currently being pursued: mechanical scanners, sensored free hand techniques, free hand without location sensing, and two dimensional arrays (Fenster & Downey, 1996).
In this scanning approach, a series of 2DUS images is recorded rapidly while the conventional transducer is manipulated over the anatomy using a variety of techniques. This digitally recorded ultrasound images are then reconstructed into 2DUS data and available for viewing. To avoid geometric distortions and inaccuracies, the relative position and angulations of each 2D image must be known accurately. One way to accomplish this is to use mechanical means to move the conventional transducer over the anatomy in a precise and predefined manner. As the transducer is moved, the 2DUS images generated by the ultrasound machine are acquired a predefined positional intervals (Gilju et al., 1994).
Then they are commonly digitized into an external computer or stored in original format in the ultrasound system computer. Then either this or that reconstructs the 3DUS data using the predefined geometric information which related the digitized 2DUS images to each other. To ensure that the scanning procedure avoids missing any region, the spacing interval between the digitized 2DUS images is precomputed and usually is made adjustable to minimize the scanning time while optimally sampling the volume (King et al., 1991).
A number of investigators and commercial companies have developed different types of mechanical assemblies used to produce 3DUS data, this can be divided in to three basic types of motions: linear, tilt, rotation (Nelson & Pretorius, 1997).
In linear scanning, the transducer translates linearly over the patient's skin so that the acquired 2DUS images are all parallel to each other. This is accomplished by mounting the conventional linear or curved transducer in an assembly housing a motor & drive mechanism such as a lead screw. When the motor is activated, the lead screw rotates, moves the transducer parallel to the skin surface.
If 2DUS images are acquired with uniform linear translation motion at regular temporal intervals then a set of parallel 2DUS images separated by regular spatial interval will be obtained, if the temporal sampling interval is fixed, then the speed of translation can be varied to change the image sampling distance of the acquisition. The ability to vary the sampling interval is important because it always allows imaging the organs in 3D with an appropriate sampling interval for the particular elevational resolution of the transducer and examination depth . Three dimensional reconstruction obtained from linear scanning has been shown to require less than 0.5 second after acquisition of 200 images (Nelson et al., 1999).
In tilt scanning, the mechanical assembly tilts the transducer about an axis parallel to the axis of the transducer. This type of motion using the external fixture approach allows the transducer face to be placed at a single location on the patient skin. Activating the motor causes the transducer to pivot on the point of contact of the skin. In this way, housing is placed against the skin while the transducer is tilted and slides against the housing to produce an angular sweep. The process of acquiring the 2D images at fixed temporal intervals during the tilting motion or at regular angular intervals results in computer storage of a set of 2D image planes arranged in a fan like geometry. This approach sweeps out a large region of interest with a fixed predefined angular separation (Nelson et al., 1999).
In rotational scanning, the ultrasound transducer is placed into an external assembly or incorporated inside a 3D mechanical probe, a motor rotates the transducer array by more than 180 degree around an axis that is perpendicular to the array and bisects it. This rotation geometry allows the axis of geometry to be fixed, while the acquired images sweeps out a conical volume in a propeller like fashion. The images intersects along the central rotation axis, with the highest spatial sampling being near the axis and poorest away from it (Nelson et al., 1999).
Figure (1): Rotational and translational scanning in 3DUS (Kurjac & Kupesic, 2000).
2-Free Hand Scanning with Position Sensing
Although the mechanical scanning approach do US of low spectral accuracy at times the bulkiness & weight of the devices hinder the scan, particularly when imaging large structures. To overcome this problem many investigators have attempted to develop various free hand scanning techniques in which the operator can hold the transducer with a position sensor attachment and manipulate it over the anatomy in the usual manner.
Over the past two decades, four basic positions sensing techniques have been developed:
C-magnetic field sensors
D-image correlation sensors
This method uses sound emitting devices such as spark gaps, mounted in transducer and array of fixed microphones which are typically mounted over the patient. During the scan ,the sound emitting devices are active and microphones continuously receive sound pulses with 2D images .the position and orientation of the recorded images are determined from knowledge of speed of sound in air and time of flight of sound pulses from the emitter to the microphones (Ohbuchi et al., 1992).
A transducer mounted on a multiple-jointed mechanical arm system. Potentiometers located at the joints of the movable arms provide the information about their relative angulations during the scan. Sufficient accuracy can be achieved by keeping individual arms, as short as possible and reducing the number of movable joints (Leotta et al., 1997).
C-Magnetic field sensors
This approach uses a transmitter which produces a spatial varying magnetic field and a small receiver containing three orthogonal rods to sense the magnetic field strength. By measuring the magnetic field strength of three components of the local magnetic field, the ultrasound transducer position and angulation can be determined (Detmer et al., 1994)
D-Image correlation techniques
The rate at which the image features change is a function of the transducer frequency, beam forming properties and patient anatomy. With careful calibration, the change in image feature can be measured (Chen et al., 1997).
3-Free Hand Scanning Without Position Sensing
In this way, the transducer is moved over the patient without any sensing device that's why the transducer must be moved in a regular and a steady manner so that the images are obtained with spacing as regular as possible (Downey &Fenster, 1995).
4-Two dimensional transducer arrays
A better approach would to keep the transducer stationary but using electronic scanning to sweep the ultrasound beam over the volume. A number of 2D arrays design have been described but the one developed at Dukes university for real time 3D echocardiography is the most advanced (Smith et al., 1992).
Figure (2): Different probe movements for 3DUS acquisition (Kurjac & Kupesic, 2000).
3D Volume Reconstruction
This process refers to the generation of a three dimensional representation of the anatomy and involves placing each acquired 2DUS image at its correct relative position to all other images. The three dimensional reconstruction process can be implemented using two distinct methods: feature based & voxel based reconstruction (Fishman et al., 1991).
A-Feature based reconstruction:
The 2DUS images are first analyzed and desired feature in each image is identified and classified. The main advantage is that the process of reducing the 3DUS data of the anatomy to a simple description of boundaries with reconstruction time and inexpensive viewing (Fishman et al., 1991).
B-Voxel based reconstruction
This accomplished by placing each pixel in the acquired 2D image in the correct location in the 3D volume matrix using its relative geometric information so as to build a voxel based volume. This approach doesn't remove any image information but rather preserves all the original information (Fishman et al., 1991).
Figure (3): Transmission of pixel to voxel in volume reconstruction of 3D images (Kurjac & Kupesic , 2000).
Visualization and Display Methods
Volume visualization difficulties
Volume visualization methods projects a three dimensional/four dimensional data set onto a dimensional image plane with the goal of gaining an understanding of the structure contained within the volumetric data. Medical volume visualization technique must offer understandable data representations, quick data manipulations and fast rendering to be useful for physicians (Wolffs, 1992).
One factor affecting image quality is speckle, which in ultrasound images often exceeds the specular echo density. The signal quality of the volume data can be improved by image compounding which occur when pixels in multiple planes are reprojected through the same voxel (Wolffs, 1992).
3DUS visualization algorithms usually treat volume data as an array of voxels. Although adjacent 2DUS images are generally closely spaced to each other; there may be gaps in the resulting volume due to under sampling. As a result some form of interpolation becomes necessary to fill in any gaps in the volume between acquired any images. Finally 3D median Gaussian filters may be used to improve the signal to noise ratio and reduce speckle prior to application of visualization algorithms (Bashford & Von Ramm, 1995).
A further difficulty with 3DUS visualization is that the volume is generally a dense object that is one where the majority of voxels contain non zero although not necessarily clinically relevant information (Kaufmen, 1996).
Volume visualization methods
Volume visualization methods include three basic approaches (Nelson &Elvins, 1993):
a- Slice projection.
b- Surface fitting.
c- Volume rendering.
a- Slice projection
Extraction of a planar image of arbitrary orientation at a particular location in a 3DUS data set utilizes standard co-ordinate transformations and rotations. Interactive display of planar slices offers the physician retrospective evaluation of anatomy particularly viewing of arbitrary planes perpendicular to the primary examination axis and other orientations not possible during data acquisition (Jones & Chen, 1995).
b- Surface fittings
Surface fit algorithms typically fit some type of planar surface primitive to values defined during the segmentary process. The surface fit approaches include contour connecting, marching cubes, marching tetra hydra, dividing cubes and others. Once the surface is defined, interactive display is typically faster than volume rendering methods because fewer data points are used, because surface fittings methods only traverse the volume once to extract the surfaces ,compared to re-evaluating the entire volume in volume rendering . After extracting the surfaces, rendering hardware & standard rendering methods can be used to render quickly the surface primitives each time the user changes a viewing or lighting parameter (Schroeder &Lorensen, 1996).
c- Volume rendering
Volume rendering methods map individual voxel directly onto the screen with using geometric primitives but require that the entire data set be sampled each time an image is rendered or re-rendered .in some situations a low resolution pass is sometimes used to create low quality images quickly for optimization of viewing position or parameter setting with a high resolution image produced immediately after values have stopped changing. The most often used volume visualization algorithm for the production of high quality images is Ray Casting method (Cabral et al., 1995).
Ray Casting (Nelson et al., 1999)
Ray casting uses the color and opacity of each voxel along a ray passing through a volume. The trajectory of the ray is defined by the relative viewing orientation of the observer and volume. Alone a particular ray, the intensity of the emerging ray is determined by the current voxel and the intensity of the entering ray.
The opacities, shades and colors encountered along the rays are blended and final opacity and color of Rout is a pixel in the image on the display.
Quantitative Analysis Methods
In 3D sonography, measuring the volume of an organ or lesion is the goal rather than measuring its diameter or perimeter. This volume measurement suffers many deficiencies (Nelson et al., 1999):
The major two deficiencies are accuracy & reproducibility.
The accuracy of a measurement technique reflects its ability to measure the truth (e.g. how close the measured volume is to the actual organ volume). Because each measurement is subject to uncertainty, the accuracy is estimated from the mean value of multiple measurements of the same structure. The variation of these measurements about the mean provide information about the precision, but the value of the mean itself determines the accuracy (Elliots et al., 1996).
The reproducibility of a measurement technique reflects its consistency (i.e. the mean variation between different measurements). This variation may be due to a number of factors and may result in a volume estimate that may have higher accuracy, average of over many measurements but that may also vary greatly among different measurements (Riccabona et al., 1996).
Volume measurement is possible by 3D & 2D also .most volume measurements made using conventional 2Dultrasound methods generally are accurate to within +/- 5% if the organs are regularly shaped (i.e. spherical) but are only accurate within +/- 20% when they are irregularly shaped. With a 3DUS image, the volume of organ can be measured either by automated technique (volume segmentation) or manually (manual plannimetry) (Brinkley et al., 1984).
The multiplanar reformatting technique is used to slice the 3D image in to a series of uniformly spaced parallel 2D images. For each 2D image, the cross sectional area of the organ is manually outlined on the computer screen using a mouse or a trackball. The areas are then summed and multiplied by the inter-slice distance to obtain an estimate of the organ volume (Nelson et al., 1999).
Volume segmentation is to use computed methods that recognizes structures automatically requiring minimal user involvement. Segmentation technique have been applied to a number of 3DUS imaging applications example ranges from simple thresolding of the image to using sophisticated boundary detection technique in cardiology to identify the ventricle.
In 3DUS successful segmentation approaches have been based on the definition and recognition of organ boundary contours. However, reliable boundary contour identification require that sufficient contrast be present in the image , where as ultrasound images uniquely suffer from different image artifacts (Baba & Jurkvic , 1997).
Sonographic Evaluation of the Normal Placenta
Two surfaces of the placenta merit special attention because they are important in assessing normal placental anatomy, evaluating placental abruption, and grading the placenta.
The fetal surface of the placenta (portion of the placenta nearest the amniotic cavity) is represented by the echogenic chorionic plate that courses along the placental tissue and is found at the junction with the amniotic fluid. This linear density is further enhanced by the strong interface of the amnion covering the chorionic plate. The second surface is the basal plate or maternal portion of the placenta, which lies at the junction of the myometrium and the substance of the placenta (Leveine &Chie 2006).
Maternal blood vessels from the endometrium (endometrial veins) run behind the basal plate and are often confused with placental abruption. This represents the normal; vascularity of this region. The endometrial veins are more apparent when the placenta is located in the fundus or posteriorly within the uterine cavity (Leveine &Chie 2006).
The substance of the placenta assumes a relatively homogenous pebble-gray appearance during the first part of pregnancy and is easily recognized with its characteristically smooth borders. The thickness of the placenta varies with gestational age, with a minimum diameter of 15 mm in fetuses greater than 23 weeks. The size of the placenta rarely exceeds 50 mm in the normal fetus (Leveine &Chie 2006).
The sonographer must maintain a perpendicular measurement of the placental width in relation to the myometrial wall when evaluating the width of the placenta. Braxton-Hicks contractions should not be confused for placental pathology.
Figure (4): Transvaginal scan at 8.5 weeks shows an early placenta (P) with chorion laeve opposite (open arrow). F, fetus; arrow, yolk sac; arrowhead, amnion ((Leveine &Chie 2006).
Several sonolucent areas within the placenta may confuse the sonographer unfamiliar with the wide range of placental variants. Cystic structures representing large fetal vessels are commonly observed coursing behind the chorionic plate and between the amnion and chorion layers. Real-time observation of blood flow or color Doppler helps to differentiate these vessels. Deposits of fibrin may also be found in the intervillous space posterior to the chorionic plate, and blood flow will not be seen in fibrinous areas (Leveine &Chie 2006).
Echo-spared regions may also be seen within the placental substance in the center of the placental lobes (cotyledons), which have been referred to as placental lakes. Blood flow should be identified within these areas. Placental veins may also be seen within the mass of the placenta (Leveine &Chie 2006).
Figure (5): Anterior placenta at 26 Wks. placenta (P) maintains a typical granular echotexture. Vessels are seen along the fetal surface, particularly adjacent to the umbilical cord insertion (arrow). Retroplacental draining veins (arrowheads) are visible. F, fetus(left). Color flow Doppler weeks demonstrates venous drainage (arrows) of the placenta (right) (Leveine &Chie 2006).
Figure (6): Intraplacental lake (left) & Color Doppler demonstrates flow within the "lake." (right) (Leveine &Chie 2006)
Placental location assessment by ultrasound (Leveine &Chie 2006)
The sonographer should always describe the position of the placenta. The placenta should be scanned longitudinally to see whether it extends into the lower segment. If it does, a transverse scan should be obtained to determine whether the placenta is located centrally or whether it lies to one side of the cervix. Oblique scans may be necessary to visualize the relationship of the placenta to the cervix.
For the sonographer to visualize the internal os of the cervix, the patient should have a full bladder. In this way the relationship of the placenta to the internal os can be visualized.
In theory this works well. However in practice, it is not always easy for the sonographer to view the internal os with the patient's bladder full. If a patient is actively bleeding or in active labor, the sonographer may not have time to wait for the patient to fill her bladder. If the fetal head is low in the pelvis, diagnosis of a posterior placenta previa may be difficult because the fetal skull bones block transmission of the ultrasound at a critical point. If the fetal head can be elevated out of the pelvis, it may be possible to distinguish between a posterior low-lying position and a posterior previa position. Other methods to demonstrate the os include tilting the patient in a slight Trendelenburg position (head lower than body) to relieve pressure of the uterus on the lower uterine segment.
An overfilled bladder may push the internal os up, making it appear higher than it actually is. This may give the false impression of a previa. Emptying the bladder reduces the pressure on the lower uterine segment and allows the cervix to assume a more normal position. The placenta may in fact not be a previa at all.
Recently vaginal ultrasound has been found helpful in cases in which the diagnosis is difficult for technical reasons, particularly with posterior placentas. This technique affords greater diagnostic accuracy in such cases compared with the traditional transabdominal approach. When performed by an experienced sonographer and in the absence of clinical bleeding. Introduction of a vaginal probe can facilitate the diagnosis. When 5-mHz or 6.5-mHz transducers are used, the focal zone is 3-12 cm or 2-7 cm, respectively. Contact with the cervix and the associated risk of hemorrhage are thus avoided.
The transperineal approach is also useful in evaluating the lower uterine segment when the definition of the placenta needs to be clarified. The transvaginal transducer is ideal for this approach. (The transducer should be prepared as for a transvaginal examination, with a protective covering). The transducer is placed along the maternal labia to demonstrate the maternal bladder, the internal os (directed in a vertical orientation), the lower uterine segment, the fetal head, and the placenta (if previa). Longitudinal and transverse scans are carefully made to delineate the relation of the placenta to the cervical os.
Accurate diagnosis has also been achieved by employing magnetic resonance imaging. This appears to improve accuracy over sonography even in cases of low posterior placental implantations. However, expense and limited availability make this a less than ideal diagnostic technique.
In 1842, Christian Johann Doppler described the Doppler effect, which describes the change in observed frequency of sound or light waves when there is relative motion between the source and the observer.
When a beam of energy waves encounters a reflector smaller than its wavelength, the incident energy waves are reflected in all directions, i.e., scattered. A portion of the scattered energy will be reflected back to the source. i.e. backscattering. In blood, the primary sources of ultrasonic scattering are the circulating red blood cells (RBC's) (Shung et al., 1976).
Theoretical and experimental evidences suggest that scattering is caused by fluctuations in red cell concentration and not by red cells acting as individual scatterers. The greater the fluctuations in red cell concentration, the more intense the scattering. The moving red cells also cause Doppler shift of the scattered US (Evans et al., 1989).
Doppler shift is a physical principle that states that when a source of sound waves is moving relative to an observer, the observer detects a shift in the wave frequency. Thus, when a sound wave strikes a moving target, the frequency of the sound waves reflected back is shifted proportionate to the velocity and direction of the moving target. Because the magnitude and direction of the frequency shift depend on the relative motion of the moving target, their velocity and direction can be determined (Maulik, 1995).
Doppler signal consists not only of blood flow generated frequency shift but also contains high-amplitude low-frequency signals, known as "clutter", produced by the movement of tissue structures and high-frequency noise generated by instrumentation. These additional signals are removed by both low and high-pass filters, respectively, in the equipment (Evans et al., 1989).
Figure (7): The Doppler effect (fd) is dependent on the velocity of flow (V) of the blood within a vessel, the initial frequency of the ultrasound beam (fc), and the cosine of the angle (A) that the ultrasound beam makes with the direction of flow. The Doppler effect is displayed on the monitor as a time-dependent plot of the frequency shift (fd) within a cardiac cycle (Leveine &Chie 2006).
Hemodynamic Information from Doppler
The Doppler power spectrum contains an immense amount of hemodynamic information from the target circulation and is usually displayed as a sonogram.
In this sonographagic display, figure (8), the vertical axis shows the magnitude of frequency shift, the horizontal axis represents the temporal change, and the brightness of the spectrum is indicative of the amplitude or the power of the spectrum (Kierney & Zimmerman, 1987).
Doppler US can generate a wide range of hemodynamic information e.g. recognition of the presence of flow, the velocity profile, quantification of flow, and assessment of downstream vascular impedance (Maulik, 1995).
Q1 = A1.V1
Where Q1 is the instantaneous flow, A1 the vascular cross-sectional area at the instant velocity measurement, and V1 the spatial mean velocity across the vascular cross-sectional area at the instant of the measurement.
Unfortunately, there are several technical difficulties with measurement of vessel diameters. The vessels are dynamic with changing diameters during the cardiac cycle, and their measurement is very sensitive to the angle of inosnation. Moreover, this methodology has high error rates ranging between 15-50% (Gill, 1985).
2- Determination of Velocity Profile:
A full spectral display is capable of providing information on the velocity profile of the target circulation. In a flat velocity profile, most RBCs travel at the same speed, this generates a slender band of Doppler frequency shift. The spectral narrowing at early systole and the relative broadening in end diastole illustrates the cardiac cycle-related changes in the velocity profile.
In a parabolic velocity profile, RBCs level at varying speeds with the cells at the center of the vessel travelling at the greatest velocity and those near the vascular wall at the least velocity.
Errors encountered in flow quantitation methodology led to development of several indirect indices of flow that compare different parts of the waveform. Hence most of Doppler indices (DIs) are independent of the angle of insonation and do not require measurement of the diameter of the vessel, they provide useful information about flow without engendering excessive errors (Hagen-Ansert, 2001). Doppler indices estimated from the maximum frequency shift envelope, figure (8).
The S/D Ratio
S/D ratio (the simplest index) is the ratio of the maximal systolic flow velocity (the peak systolic frequency shift, S) to the minimal end-diastolic flow velocity (the end-diastolic frequency shift, D).
During pregnancy, the uterine and umbilical arteries normally maintain diastolic flow and the normal placental bed are characterized by low resistance and high blood flow. Thus the most useful S/D ratios are obtained from the maternal uterine and fetal umbilical arteries and provide an indirect estimate of the adequacy of blood flow to the fetus. Because low diastolic velocities are seen in more central fetal vessels, e.g. the descending aorta, the S/D ratio is not useful elsewhere in the fetal circulation (Low, 1991).
Figure (8): Doppler indices estimated from the maximum frequency shift (MFS) envelope (S =Systole / D = Diastole / A = Temporal average frequency shift in one cardiac cycle. (coted from Doppler ultrasound in obstetric &gynecology by Dev Maulik)
The D/A Ratio
Because the variations in the end-diastolic frequency shift appear to be the more relevant component of the waveform, Maulik et al., 1984 suggested the use of this parameter (D) normalized by the mean value of the maximum frequency shift envelope over the cardiac cycle.
The Resistance Index (RI)
RI = (S-D) / (S)
Where (S) peak systolic flow and (D) peak diastolic flow.
This ratio is applicable only to the umbilical and the uterine arteries, as low diastolic values limit the usefulness in the fetal aorta or other central vessels (Yarlagadda et al., 1989).
The Peak-To-Peak Pulsatility Index (PI)
PI=(S-D) / (A)
Where (S) peak systolic frequency shift, (D) the end-diastolic frequency shift and (A) the time average velocity (A).
PI requires a digitized waveform for calculating the mean of the maximal frequencies represented. Because of the mean value in the denominator, this index can be computed using flow data from the fetal descending aorta without encountering the excessive variation that can be caused by small numbers as with S/D and RI indices (Maulik, 1995).
The Notch Index (NI)
NI = (C-D) / C
Where (D) post-systolic nadir and (C) following length of the waveform.
To find a simple discrimination between normal and pathological uterine perfusion, best diagnostic performance was reached by a definition using a combination of high impendance and notch:
No notch and mean PI > 95th centile.
Unilateral notch and mean PI > 90th centile.
Bilateral notch and mean PI > 50th centile.
The prevalence of notch in nulliparae (8.5%) was higher than in parae (4.7%) and decreased with increasing gestational age (20 weeks: 8.6% - 23 weeks 5.4%). We found a clear relation between elevation of Impedance, depth of notch and frequency of adverse pregnancy outcome with a frequency of complications varying from 3.2% (mean PI < 0.8, mean NI = 0.1) to 38.4% (mean PI > 2.0, mean NI>0.1) (Becker et al., 2002).
Uterine Artery Doppler
Uterine artery can be examined using either CW or PW Doppler. the probe is directed into the parauterine area in the region of the lower uterine segment. Insonation of the uterine artery must be at its crossover the iliac artery, however the external iliac artery can be distinguished from uterine arteries because iliac arteries normally do not have diastolic flow (Frusca et al., 1997).
Uterine artery Doppler waveform shape is unique, characterized by high diastolic velocity similar to those in systole and highly turbulent flow, with many different velocities apparent. The diastolic velocity increases and thus the indices decrease as term approaches (Albaiges et al., 2000). A failure of this pattern to appear or the presence of a notch in the waveform at end-systole has been reported with IUGR (Harrington et al., 1996).
In normal pregnancy, impedance to flow in uterine arteries decrease with gestation (figure 9&10), which may be the consequences of trophoplastic invasion of spiral arteries and their conversion into low resistance vessels. Pre-eclampsia and IUGR are associated with failure of trophoplastic invasion of spiral arteries. Doppler studies, in these conditions, have shown that impedance to flow in the uterine arteries is increased (Irion et al., 1998).
Figure (9): Flow velocity waveform obtained from uteroplacental circulation at 8 &16 wks gestation. Note progressive increase in diastolic flow in uteroplacental circulations as gestational age increases (Maulik, 2005).