阅读说明：本技术 用于执行脉搏波速度测量的系统和方法 (System and method for performing pulse wave velocity measurements ) 是由 M·P·J·屈嫩 A·范德霍斯特 于 2019-08-13 设计创作，主要内容包括：本发明提供用于基于沿着血管引导的多个血管内超声脉冲来计算脉搏波速度的系统和方法。针对每个超声脉冲,多个回波从沿着所述血管的多个距离接收。第一超声多普勒信号从距脉冲原点的第一距离接收,并且第二超声多普勒信号从距脉冲原点的第二距离接收。第一和第二流速度量分别地基于第一和第二超声多普勒信号获得。所述脉搏波速度基于时间延迟来计算,所述时间延迟基于所述第一流速度量和所述第二流速度量。(The present invention provides systems and methods for calculating pulse wave velocity based on a plurality of intravascular ultrasound pulses directed along a blood vessel. For each ultrasound pulse, a plurality of echoes is received from a plurality of distances along the blood vessel. The first ultrasonic doppler signal is received from a first distance from the pulse origin and the second ultrasonic doppler signal is received from a second distance from the pulse origin. The first and second flow velocity quantities are obtained based on the first and second ultrasonic doppler signals, respectively. The pulse wave velocity is calculated based on a time delay based on the first and second flow rate metrics.)

1. A system (2) for performing intravascular pulse wave velocity measurements, the system comprising:

an intravascular ultrasound device (200) adapted to:

generating a plurality of intravascular ultrasound pulses (230) directed along a blood vessel (220) at a pulse origin;

receiving, for each ultrasound pulse, a plurality of echoes (240, 250) from a plurality of distances along the blood vessel; and

a processor (28) adapted to:

obtaining a first distance (z) from the origin of the pulse₁) And from a second distance (z) from the origin of the pulse₂) The second ultrasonic doppler signal of (a);

determining a time delay (Δ t) in the velocity characteristic (270) at the first distance to modify to propagate to the second distance based on the obtained first and second ultrasonic Doppler signals;

calculating the pulse wave velocity based on the time delay.

2. The system of claim 1, wherein the intravascular ultrasound device comprises:

a single ultrasound transducer element; or

An array of ultrasound transducer elements.

3. The system of claim 1 or 2, wherein determining the time delay comprises performing a cross-correlation between the first ultrasonic doppler signal and the second ultrasonic doppler signal.

4. The system of claim 1 or 2, wherein determining the time delay comprises:

obtaining a first flow velocity metric and a second flow velocity metric based on the first ultrasonic Doppler signal and the second ultrasonic Doppler signal.

5. The system of claim 4, wherein the first flow rate amount comprises a first average speed and the second flow rate amount comprises a second average speed.

6. The system of claim 4, wherein the first flow rate quantity comprises a first profile of flow rates and the second flow rate quantity comprises a second profile of flow rates.

7. The system of claim 6, wherein the first distribution of flow rates comprises a first frequency spectrum and the second distribution of flow rates comprises a second frequency spectrum.

8. The system of claim 6 or 7, wherein determining the time delay comprises:

calculating an individual time delay for each velocity value in the first and second distributions of flow velocities; and is

Calculating an average time delay based on the individual time delays.

9. The system of claim 6 or 7, wherein determining the time delay comprises:

extracting a first feature of the first distribution of flow rates;

extracting a second characteristic of the second distribution of flow velocities; and is

Calculating a time delay based on the first feature and the second feature.

10. The system of claim 9, wherein the first and second features comprise one or more of:

instantaneous peak velocity;

starting a pulse;

a peak time;

a maximum acceleration; and

the instantaneous average speed.

11. The system of claim 6 or 7, wherein determining the time delay comprises performing a cross-correlation between a first distribution of the flow rates and a second distribution of the flow rates.

12. The system of any one of the preceding claims, wherein the system is further configured to obtain a pressure metric, and wherein the calculation of the pulse wave velocity is based on the time delay and the pressure metric.

13. The system of claim 1, wherein the guidance of the plurality of intravascular ultrasound pulses along the blood vessel comprises electronic beam steering and/or electronic beam focusing.

14. A method (100) of calculating pulse wave velocity, the method comprising:

generating a plurality of intravascular ultrasound pulses (110) guided along a blood vessel at a pulse origin;

receiving, for each ultrasound pulse, a plurality of echoes (120) from a plurality of distances along the blood vessel;

obtaining a first ultrasonic Doppler signal (130) from a first distance from the pulse origin and a second ultrasonic Doppler signal (135) from a second distance from the pulse origin;

determining (150), based on the obtained first and second ultrasonic Doppler signals, a time delay for a modification in the velocity characteristic at the first distance to propagate to the second distance;

calculating (160) the pulse wave velocity based on the time delay.

15. A computer program comprising computer program code means adapted to implement the method of claim 14 on a system according to any one of claims 1 to 13 when the computer program is run on a computer.

Technical Field

The present invention relates to a system for measuring pulse wave velocity, and in particular to an intravascular system for measuring pulse wave velocity.

Background

Arteriosclerosis is an important risk factor for cardiovascular diseases. Arterial stiffness increases with aging and various disease states, including: hypertension; high cholesterol; diabetes mellitus; obesity; smoking; and renal diseases.

A common parameter for assessing the stiffness of a blood vessel is the Pulse Wave Velocity (PWV). PWV is the transmission speed of pressure/flow waves generated by the heart beat, for example, through an artery. PWV is determined by the ability of the vessel to dilate (i.e., distensibility D) according to the following relationship:

wherein the content of the first and second substances,wherein; v is the blood vessel volume; p is the intravascular pressure; and ρ is the blood density. From this relationship, the Moens-Korteweg equation can be derived:

wherein: e is Young's modulus; d is the vessel diameter; and h is wall thickness. By evaluating PWV, the stiffness of the artery can be quantified. A typical value for PWV in an artery is 10m/s, which is an order of magnitude higher than the velocity of blood particles.

The relevance of local PWV in arteries in pilot treatment is clear from recent studies, such as: systematic evaluation of halmokinetic parameters to predicted halmokinetic responses to residual identity preservation (Abstract EuroPCR, 2016) by Finegold et al. The article indicates that PWV within the main renal artery pretreatment predicts the outcome of renal denervation in patients with resistant hypertension. A Development of Coronary Pulse Wave Velocity in Harbaoui et al, New Pathophysiological Insight Intra Coronary Artery Disease (J Am Heart Assic.2017) indicates that low Coronary PWV is associated with acute Coronary syndrome, indicating a relationship between plaque vulnerability and arterial stiffness.

When pressure/flow waves traverse blood vessels very rapidly, PWVs are most commonly determined over relatively large distances in the vasculature, such as from the brachial artery to the ankle joint. In this way, the average stiffness of the blood vessels between the measurement locations is determined.

Over the past decade, additional local PWV measurements have been developed. For superficial arteries, external ultrasound may be used to assess PWV. Another method that is also suitable for local assessment of PWV in deeper arteries is the use of sensor-equipped catheters. For example, two or more pressure sensors on a catheter at a known distance (x) may be used to determine the time difference (Δ t) of the passing wave as:

currently, these catheter-based measurements are typically performed in the aorta, which is relatively long and has a large diameter. For shorter arteries, which are also typically much smaller in diameter, the technical requirements of the measurement device (such as sampling frequency, synchronization, etc.) are more challenging.

Furthermore, all proposed solutions require at least two sensors (such as dual pressure sensors, or pressure and flow sensors). This increases the complexity and cost of the PWV measurement device.

Therefore, there is a need for a PWV measurement system that can perform PWV measurements in smaller arteries without requiring significant additional hardware.

Disclosure of Invention

The invention is defined by the claims.

According to an example of an aspect of the present invention, there is provided a method for calculating a pulse wave velocity, the method including:

generating a plurality of intravascular ultrasound pulses directed along a central axis of the blood vessel at a pulse origin;

receiving, for each ultrasound pulse, a plurality of echoes from a plurality of distances along a central axis of the blood vessel;

obtaining a first ultrasonic doppler signal from a first distance from the pulse origin and a second ultrasonic doppler signal from a second distance from the pulse origin;

calculating a time delay based on the first ultrasonic Doppler signal and the second ultrasonic Doppler signal; and is

Calculating the pulse wave velocity based on the time delay.

The method provides for calculating the pulse wave velocity in a blood vessel using a single sensor, such as an ultrasound transducer.

In pulse doppler processing, doppler signals are obtained from received echo signals at various measurement depths, which correspond to certain time delays after each pulse transmission. The measurement depth and time delay are linked by the known speed of sound in the blood. Each transmission will provide one sample to the doppler signal at each measurement depth.

Since the first and second measurement distances are known and can be predetermined based on system constraints to result in the highest possible signal quality, the time delay associated with the received doppler signal between the two measurement locations can be used to calculate the pulse wave velocity.

In other words, the doppler signal in the (blood) vessel is dynamically measured in at least two positions along the main axis of the vessel of interest. The pulse wave velocity is then calculated from the time it takes the doppler signal profile to travel from the first location to a second (or other) location. An advantage is that the velocity signal profiles need not be the same (e.g. in terms of amplitude, width) as it is sufficient, if the velocity profiles show similarities, as for the PWV calculation only the time interval is necessary for the velocity profiles to travel from the first position to the second position. In other words, due to a misalignment of the intravascular device relative to the longitudinal axis of the blood vessel, the velocity profile measured at a first location may be truncated (measured) at a second location having a significantly different profile. This means that misalignment of the intravascular device with respect to the longitudinal axis of the blood vessel is tolerated as long as two positions of velocity measurements occur within the lumen of the blood vessel.

In an embodiment, the calculation of the time delay comprises performing a cross-correlation between the first and second ultrasonic doppler signals.

In this way, the time delay may be determined by acquiring the sample time delay between the first and second measurement locations at which the doppler signal provides the highest correlation.

In an embodiment, the calculation of the time delay comprises:

obtaining a first flow velocity metric based on the first ultrasonic Doppler signal;

obtaining a second flow velocity metric based on the second ultrasonic Doppler signal; and is

Calculating the time delay based on the first flow rate metric and the second flow rate quantity.

In this way, it is possible to calculate the time delay based on the first and second flow velocity amounts derived from the first and second ultrasound doppler signals, respectively.

In an arrangement, the first flow rate amount comprises a first average speed and the second flow rate amount comprises a second average speed.

The velocity at the measurement location may be averaged over time in order to improve the accuracy of the time delay calculation.

In another arrangement, the first flow rate quantity comprises a first distribution of flow rates and the second flow rate quantity comprises a second distribution of flow rates.

The flow velocity profile provides a full range of velocities measured at measurement locations on the plurality of shots, thereby providing a full representation of the vascular measurement.

In another arrangement, the obtaining of the first distribution of flow velocities includes obtaining a first frequency spectrum, and the obtaining of the second distribution of flow velocities includes obtaining a second frequency spectrum.

The flow velocity profile is typically collected as a frequency spectrum. This may allow analysis to be performed across frequency bins, which would be equivalent to a given speed value or range of values. The size of the frequency bins may vary based on system limitations.

In another or other embodiments, the calculating of the time delay includes:

calculating an individual time delay for each velocity value of the first distribution of flow velocities and the second distribution of flow velocities; and is

Calculating an average time delay based on the individual time delays.

In this way, each measurement speed can be considered. By averaging the velocities over the distribution, it is possible to reduce the impact of erroneous measurement results on the accuracy of the final result.

In another or other embodiments, the calculating of the time delay includes:

extracting a first feature of the first distribution of flow rates;

extracting a second characteristic of the second distribution of flow rates; and is

Calculating a time delay based on the first feature and the second feature.

In this way, aspects of the velocity profile that do not correspond to blood flow (such as slow velocities representing wall motion rather than flow) may be removed from the calculation, thereby increasing the accuracy of the final result.

In another embodiment, the first feature is a first plurality of features and the second feature is a second plurality of features.

By taking into account a number of features, the accuracy of the final calculation is increased.

In yet another embodiment, the first and second features include one or more of:

instantaneous peak velocity;

starting a pulse;

a peak time;

a maximum acceleration; and

the instantaneous average speed.

In an embodiment, the calculation of the time delay comprises performing a cross-correlation between a first distribution of flow rates and a second distribution of flow rates.

In an embodiment, the method further comprises obtaining a pressure measure, and wherein the calculation of the pulse wave velocity is based on the time delay and the pressure measure.

By considering a plurality of methods of calculating the pulse wave velocity, the accuracy of the final calculation can be increased.

In an arrangement, the directing of the plurality of intravascular ultrasound pulses along the central axis of the blood vessel comprises electronic beam steering and/or electronic beam focusing.

According to an example of an aspect of the present invention, there is provided a computer program comprising computer program code means adapted to perform the above described method when said computer program is run on a computer.

According to an example of an aspect of the present invention, there is provided an ultrasound system for performing intravascular pulse wave velocity measurements, the system comprising:

an intravascular ultrasound unit adapted to:

generating a plurality of intravascular ultrasound pulses directed along a central axis of the blood vessel at a pulse origin;

receiving, for each ultrasound pulse, a plurality of echoes from a plurality of distances along a central axis of the blood vessel; and

a processor adapted to:

obtaining a first ultrasonic doppler signal from a first distance from the pulse origin and a second ultrasonic doppler signal from a second distance from the pulse origin;

calculating a time delay based on the first ultrasonic Doppler signal and the second ultrasonic Doppler signal; and is

Calculating the pulse wave velocity based on the time delay.

In an embodiment, the intravascular ultrasound unit comprises:

a single ultrasound transducer element;

an array of ultrasound transducer elements; or

A first intravascular ultrasound element and a second intravascular ultrasound element, wherein the first intravascular ultrasound element and the second intravascular ultrasound element are spatially independent of each other.

These and other aspects of the invention are apparent from and will be elucidated with reference to the embodiment(s) described hereinafter.

Drawings

Examples of the invention will now be described in detail with reference to the accompanying drawings, in which:

FIG. 1 shows an ultrasonic diagnostic imaging system explaining the general operation;

FIG. 2 illustrates the method of the present invention;

fig. 3 shows a schematic representation of an intravascular device;

FIG. 4 shows a plot of velocity versus time for first and second measurement locations;

figure 5 shows a series of graphs relating to a plurality of ultrasonic doppler signals obtained at various depths;

FIG. 6 shows a time window of the graph of FIG. 5;

FIG. 7 shows a graphical representation of estimated Loupas velocity;

FIG. 8 shows a graph of cross-covariance for velocity signals obtained at different measurement depths; and is

Fig. 9 shows a graph of measured depth and maximum cross-covariance.

Detailed Description

The present invention will be described with reference to the accompanying drawings.

It should be understood that the detailed description and specific examples, while indicating exemplary embodiments of the devices, systems, and methods, are intended for purposes of illustration only and are not intended to limit the scope of the invention. These and other features, aspects, and advantages of the apparatus, systems, and methods of the present invention will become better understood with regard to the following description, appended claims, and accompanying drawings. It should be understood that the figures are merely schematic and are not drawn to scale. It should also be understood that the same reference numerals are used throughout the figures to indicate the same or similar parts.

The present invention provides a method for calculating a pulse wave velocity based on a plurality of intravascular ultrasound pulses directed along a central axis of a blood vessel. For each ultrasound pulse, a plurality of echoes are received from a plurality of distances along a central axis of the blood vessel. The first ultrasonic doppler signal is received from a first distance from the pulse origin and the second ultrasonic doppler signal is received from a second distance from the pulse origin. The first and second flow velocity quantities are obtained based on the first and second ultrasonic doppler signals, respectively. The pulse wave velocity is calculated based on a time delay, which is based on the first flow rate metric and the second flow rate metric.

The general operation of an exemplary ultrasound system will first be described with reference to figure 1 and the emphasis on the signal processing functions of the system, as the present invention relates to the processing of signals measured by a transducer array by the system.

The system includes an array transducer probe 4 having a transducer array 6 for transmitting ultrasound waves and receiving echo information. The transducer array 6 may comprise CMUT transducers; piezoelectric transducers formed of materials such as PZT or PVDF; or any other suitable transducer technology. In this example, the transducer array 6 is a two-dimensional array of transducers 8 capable of scanning a 2D plane or three-dimensional volume of the region of interest. In another example, the transducer array may be a 1D array.

The transducer array 6 is coupled to a microbeamformer 12, the microbeamformer 12 controlling the reception of signals by the transducer elements. The microbeamformer is capable of at least partially beamforming signals received by sub-arrays (often referred to as "groups" or "tiles") of transducers, as described in U.S. patents US 5997479(Savord et al), US 6013032(Savord) and US 6623432(Powers et al).

It should be noted that the microbeamformer is entirely optional. In addition, the system includes a transmit/receive (T/R) switch 16 to which the microbeamformer 12 can be coupled and which switches the array between transmit and receive modes and protects the main beamformer 20 from high energy transmit signals without the use of a microbeamformer and with the transducer array being directly operated by the main system beamformer. The transmission of ultrasound beams from the transducer array 6 is directed by a transducer controller 18 coupled to the microbeamformer through a T/R switch 16 and a main transmit beamformer (not shown) which may receive input from user operation of a user interface or control panel 38. The controller 18 may comprise transmit circuitry arranged to drive (directly or via the microbeamformer) the transducer elements of the array 6 during a transmit mode.

In a typical line-by-line imaging sequence, the beamforming system within the probe may operate as follows. During transmit, the beamformer (which may be a microbeamformer or a main system beamformer depending on the implementation) activates the transducer array or sub-apertures of the transducer array. The sub-apertures may be one-dimensional lines of transducers or two-dimensional patches of transducers within a larger array. In transmit mode, the focusing and steering of the ultrasound beams generated by the array or sub-apertures of the array is controlled, as described below.

After receiving the backscattered echo signals from the object, the received signals undergo receive beamforming (as described below) in order to align the received signals and, in case of using a sub-aperture, then the sub-aperture is shifted, for example by one transducer element. The displaced sub-apertures are then activated and the process repeats until all of the transducer elements of the transducer array have been activated.

For each line (or sub-aperture), the total receive signal of the associated line used to form the final ultrasound image will be the sum of the voltage signals measured by the transducer elements of the given sub-aperture during the receive period. After the following beamforming process, the resulting line signals are commonly referred to as Radio Frequency (RF) data. Each line signal (RF data set) generated by the respective sub-aperture then undergoes additional processing to generate a line of the final ultrasound image. The change in amplitude of the line signal over time will contribute to the change in brightness of the ultrasound image over depth, where a high amplitude peak will correspond to a bright pixel (or set of pixels) in the final image. Peaks that occur near the beginning of the line signal will represent echoes from shallow structures, while peaks that gradually occur later in the line signal will represent echoes from structures at increasing depths within the object.

One of the functions controlled by the transducer controller 18 is beam steering and direction of focusing. The beams may be steered straight ahead from the transducer array (orthogonal thereto), or at different angles for a wider field of view. Steering and focusing of the transmit beams may be controlled as a function of transducer element actuation times.

Two methods can be distinguished in general ultrasound data acquisition: plane wave imaging and "beam steering" imaging. The two methods are distinguished by the presence of beamforming in the transmit mode ("beam steering" imaging) and/or the receive mode (plane wave imaging and "beam steering" imaging).

Looking first at the focusing function, by activating all transducer elements simultaneously, the transducer array generates a plane wave that diverges as it travels through the object. In this case, the beam of the ultrasonic wave remains unfocused. By introducing a location dependent time delay to the activation of the transducer, the wavefront of the beam can be focused at a desired point, referred to as the focal zone. The focal zone is defined as a point having a lateral beamwidth less than half the transmit beamwidth. In this way, the lateral resolution of the final ultrasound image is improved.

For example, if the time delays cause the transducer elements to activate in the series starting from the outermost element and ending at the central element(s) of the transducer array, a focal zone will be formed at a given distance from the probe, coincident with the central element(s). The distance of the focal zone from the probe will vary depending on the time delay between each subsequent wheel of transducer element activation. After the beam passes through the focal zone, it will begin to diverge, forming a far field imaging region. It should be noted that for focal zones located close to the transducer array, the ultrasound beam will diverge rapidly in the far field, resulting in beam width artifacts in the final image. In general, the near field between the transducer array and the focal zone shows little detail due to the large overlap in the ultrasound beams. Thus, changing the location of the focal region can result in significant changes in the quality of the final image.

It should be noted that in transmit mode, only one focal point may be defined unless the ultrasound image is divided into multiple focal regions (each of which may have a different transmit focal point).

In addition, after receiving the echo signal from within the object, the reverse process of the above process can be performed to perform the reception focusing. In other words, the incoming signals may be received by the transducer elements and subjected to an electronic time delay before being passed into the system for signal processing. The simplest example of this is known as delay-and-sum (delay-and-sum) beamforming. The receive focus of the transducer array can be dynamically adjusted as a function of time.

Turning now to the function of beam steering, by correctly applying time delays to the transducer elements, a desired angle can be imparted on the ultrasound beam as it leaves the transducer array. For example, by activating the transducers on a first side of the transducer array in a sequence ending on the opposite side of the array, followed by the remaining transducers, the wave front of the beam will be angled towards the second side. The magnitude of the steering angle relative to the normal of the transducer array depends on the magnitude of the time delay between subsequent transducer element activations.

Furthermore, the steered beam can be focused, wherein the total time delay applied to each transducer element is the sum of both the focusing and steering time delays. In this case, the transducer array is referred to as a phased array.

In the case of a CMUT transducer that requires a DC bias voltage for its activation, the transducer controller 18 may be coupled to control the DC bias control 45 of the transducer array. The DC bias control 45 sets the DC bias voltage(s) applied to the CMUT transducer elements.

For each transducer element of the transducer array, an analog ultrasound signal, commonly referred to as channel data, enters the system through a receive channel. In the receive channels, partial beamformed signals are generated from the channel data by the microbeamformer 12 and then passed to the main receive beamformer 20 where the partial beamformed signals from the individual patches of transducers are combined into fully beamformed signals, referred to as Radio Frequency (RF) data. The beamforming performed at each stage may be performed as described above, or may include additional functionality. For example, the main beamformer 20 may have 128 channels, each receiving partially beamformed signals from a tile of tens or hundreds of transducer elements. In this way, the signals received by the thousands of transducers of the transducer array can effectively contribute to a single beamformed signal.

The beamformed receive signals are coupled to a signal processor 22. The signal processor 22 can process the received echo signals in various ways, e.g., band pass filtering; selecting; separating I and Q components; and harmonic signal separation for separating linear and nonlinear signals, thereby enabling identification of nonlinear (higher harmonics of the fundamental frequency) echo signals returned from tissue and microbubbles. The signal processor 22 may also perform additional signal enhancement, such as speckle reduction, signal synthesis, and noise cancellation. The band pass filter in the signal processor may be a tracking filter whose pass band slides from a higher frequency band to a lower frequency band as echo signals are received from increasing depths, thereby rejecting noise at higher frequencies from greater depths, which are typically devoid of anatomical information.

The beamformers for transmission and for reception are implemented in different hardware and may have different functions. Of course, the receiver beamformer is designed to take into account the characteristics of the transmit beamformer. For simplicity, only the receiver beamformers 12, 20 are shown in fig. 1. In the overall system there will also be a transmit chain with a transmit microbeamformer and a main transmit beamformer.

The function of the microbeamformer 12 is to provide an initial combination of signals in order to reduce the number of analog signal paths. This is typically performed in the analog domain.

The final beamforming is done in the main beamformer 20 and is usually done after digitization.

The transmit and receive channels use the same transducer array 6 with a fixed frequency band. However, the bandwidth occupied by the transmit pulses may vary depending on the transmit beamforming used. The receive channel can capture the entire transducer bandwidth (which is the classical approach) or, by using band-pass processing, it can only extract the bandwidth containing the desired information (e.g., the harmonics of the dominant harmonics).

The RF signal may then be coupled to a B-mode (i.e., brightness mode or 2D imaging mode) processor 26 and a doppler processor 28. The B-mode processor 26 performs amplitude detection on the received ultrasound signals to image structures in the body, such as organ tissue and blood vessels. In the case of progressive imaging, each line (beam) is represented by an associated RF signal, the amplitude of which is used to generate the luminance values to be assigned to the pixels in the B-mode image. The exact location of a pixel within an image is determined by the location of the associated amplitude measurement along the RF signal and the number of lines (beams) of the RF signal. The B-mode image of such a structure may be formed in harmonic or fundamental image modes or a combination of both, as described in US6283919 (roudhill et al) and US 6458083(Jago et al). A doppler processor 28 processes temporally different signals produced by tissue movement and blood flow to detect moving matter, such as a flow of blood cells in an image field. The doppler processor 28 typically includes a wall filter having parameters set to pass or reject echoes returned from selected types of materials within the body.

The structural and motion signals produced by the B mode and doppler processors are coupled to the scan converter 32 and the multiplanar reformatter 44. The scan converter 32 arranges the echo signals in a spatial relationship according to which the echo signals are received in a desired image format. In other words, the scan converter is used to convert the RF data from a cylindrical coordinate system to a Cartesian coordinate system suitable for displaying ultrasound images on the image display 40. In the case of B-mode imaging, the brightness of a pixel at a given coordinate is proportional to the amplitude of the RF signal received from that location. For example, the scan converter may arrange the echo signals into a two-dimensional (2D) fan format or a pyramidal three-dimensional (3D) image. The scan converter may superimpose colors corresponding to motion at points in the image field onto the B-mode structural image, where doppler estimated velocity yields a given color. The combined B-mode structural image and color doppler image depict the motion of tissue and blood flow within the structural image field. As described in US patent US 6443896(Detmer), a multiplanar reformatter converts echoes received from points in a common plane in a volumetric region of the body into an ultrasound image of that plane. The volume renderer 42 converts the echo signals of the 3D data set into a projected 3D image as seen from a given reference point, as described in US 6530885(Entrekin et al).

The 2D or 3D images are coupled from the scan converter 32, the multi-plane reformatter 44 and the volume renderer 42 to the image processor 30 for further enhancement, buffering and temporary storage for display on the image display 40. The imaging processor may be adapted to remove certain imaging artifacts from the final ultrasound image, such as: acoustic shadowing, for example caused by strong attenuators or refraction; post-enhancement, for example caused by weak attenuators; reverberation artifacts, for example, where highly reflective tissue interfaces are located in close proximity; and so on. In addition, the image processor may be adapted to process certain speckle reduction functions in order to improve the contrast of the final ultrasound image.

In addition to being used for imaging, the blood flow values produced by the doppler processor 28 and the tissue structure information produced by the B-mode processor 26 are also coupled to a quantification processor 34. The quantification processor produces measures of different flow conditions, for example the volumetric rate of blood flow in addition to structural measurements such as organ size and gestational age. The quantification processor may receive input from the user control panel 38, such as points in the anatomy of the image at which measurements are to be taken.

The output data from the quantization processor is coupled to a graphics processor 36 for rendering the measurement graphics and values together with the image on a display 40 and for audio output from the display device 40. The graphics processor 36 may also generate a graphical overlay for display with the ultrasound images. These graphic overlays may contain standard identifying information such as patient name, date and time of the image, imaging parameters, and the like. For these purposes, the graphics processor receives input from the user interface 38, such as a patient name. The user interface is also coupled to the transmit controller 18 to control the generation of ultrasound signals from the transducer array 6, and hence the images produced by the transducer array and ultrasound system. The transmit control function of the controller 18 is only one of the functions performed. The controller 18 also takes into account the mode of operation (given by the user) and the corresponding required transmitter configuration and bandpass configuration in the receiver analog-to-digital converter. The controller 18 may be a state machine having a fixed state.

The user interface is further coupled to a multi-plane reformatter 44 for selecting and controlling a plane of a plurality of multi-plane reformatted (MPR) images, which may be used to perform a measure of quantization in the image field of the MPR images.

Fig. 2 shows a method 100 for calculating pulse wave velocity.

In step 110, a plurality of intravascular ultrasound pulses are generated at a pulse origin, which is directed along a central axis of the blood vessel.

The blood flow velocity measurement may be performed by a pulse wave doppler ultrasound method, which allows the location of interest to be measured with a single ultrasound transmission and thus by only a single sensor. This requires that the ultrasound transducer be aligned with the main axis of the blood vessel, which can usually be best achieved by means of intravascular methods.

In other words, a single ultrasound transducer may be inserted into a blood vessel to perform a blood flow measurement. This may be performed using a guide wire or catheter with an ultrasound transducer at its tip aimed along the main axis of the blood vessel.

In step 120, for each ultrasound pulse, a plurality of echoes is received from a plurality of distances along a central axis of the blood vessel.

In step 130, a first ultrasonic doppler signal is obtained from a first distance from the pulse origin, and in step 135, a second ultrasonic doppler signal is obtained from a second distance from the pulse origin.

The first and second doppler signals are constructed from ultrasound echoes received from the first and second measured ranges, respectively.

Maximum pulse wave velocity PWV in the context of pulse wave velocity estimation_maxGoverning a minimum time delay at_minIt needs to be resolved by the system to provide an accurate estimate:

wherein: z is a radical of₁Is a first distance from the origin of the ultrasonic pulse; and z is₂Is a second distance from the origin of the ultrasonic pulse. In other words, z₂-z₁Is the distance between the first and second measurement points along the blood vessel.

To achieve good resolution, Δ t_minSeveral pulses must be transmitted across the ultrasound transducer so that:

Δt_min＞＞1/f_PRF

wherein f is_PRFIs the pulse rate frequency of the ultrasound transducer, i.e. the number of ultrasound pulses generated per second.

In the example of renal denervation, pulse wave velocities are expected to be up to 20 m/s. Take this as PWV_maxAnd select z₁4mm and z₂8mm results in Δ t_min>0.2ms, which would mean an example f where 16 ultrasound pulses would be at 80kHz_PRFAnd (4) generating. Thus, it is expected that there will be a worth of velocity wave hits z₁When and whenIts hit z₂At least 16 consecutive ultrasound pulses of a delay between times are transmitted. It should be noted that z is selected₁Too small may cause the received signal, and hence the velocity metric, to be affected by the presence of the intravascular device itself. Thus, z₁Should be chosen to give sufficient distance between the device and the measurement point to reduce or eliminate such interference.

In pulse Doppler processing, the Doppler signal passes through a respective measurement depth z corresponding to some time delay after each ultrasonic pulse is transmitted_mPhase-quadrature demodulation of the detected pulse-echo signals or Radio Frequency (RF) data obtained at the receiving station.

Phase-quadrature demodulation (or complex demodulation) is equivalent to the separation of the in-phase (I) and quadrature (Q) components of the received signal in order to demodulate it from the high frequency band to the baseband. This function may be performed by the doppler processor 28 described above with reference to figure 1. Phase-quadrature demodulation is further described in Loupas et al, "An Axial vector Estimator for ultrasonic Flow Imaging, Based on a Full Evaluation of the Doppler Evaluation by Means of a Two-Dimensional automatic correlation application" (IEEE Transactions on Ultrasonics, ferroelectronics, and Frequency Control, Vol. 42, Vol. 4, 7 months 1995).

The measurement depth and time delay are related to each other by the known speed of sound in the medium being imaged (such blood in a blood vessel). Each ultrasound pulse transmission will provide one sample to the doppler signal at each measurement depth.

In step 140, a first flow velocity volume is obtained based on the first ultrasonic doppler signal, and in step 145, a second flow velocity volume is obtained based on the second ultrasonic doppler signal.

In doppler ultrasound imaging, the received signal is used to determine the velocity of a given fluid. When a signal is reflected from a fluid particle, it will receive a doppler shift according to the direction of flow of the fluid, which can then be used to determine the velocity of the fluid at the location of the received echo.

In case an intravascular doppler measurement is performed, the velocity of the blood will not be uniform and constant, but will have a time variation according to various physiological processes, such as the beating of the heart. This will cause the velocity of the blood to change in some way over a given period of time, such as within a few seconds. This characteristic change in velocity will propagate through the vascular system at a certain velocity, i.e., PWV.

By identifying this characteristic change in velocity at two known measurement locations, and determining the time it takes to occur at the second location after occurring at the first location, it is possible to measure PWV.

The characteristic change in speed may be identified by a plurality of speed metrics. The speed metric measured at the first measurement location is referred to as a first speed metric and the speed metric measured at the second measurement location is referred to as a second speed metric; however, the first and second speed metrics may represent the same characteristic change in speed.

In an example, the first flow rate amount may be a first average speed and the second flow rate amount may be a second average speed. More particularly, the flow rate may be measured at the first and second measurement points over a predetermined length of time. When, for example, an average speed as measured within 1ms occurs at the second measurement location after having occurred at the first measurement location (meaning that the first and second speed metrics match), it may be determined that a change in the characteristic of the speed has traversed the measurement distance.

The average velocities at the first and second measurement points may be evaluated, for example, by the Kasai algorithm (as described in "Real-Time Two-Dimensional Blood Flow Imaging Using An Autocorrelation Technique" of Kasai et al (IEEE Transactions on Sound and Ultrasonics, Vol. SU-32, No. 3, 5 198)) or the Loupas algorithm (as described in "An Axial vector Estimator for ultrasonic Blood Flow Imaging, Based on a Full Evaluation of the Doppler by Means of a Two-Dimensional Autocorrelation analysis, Vol. 4, 1995) or the Loupas algorithm (IEEE Transactions on Ultrasition, Ferror Freund, Ferror, Control, Vol. 4, 1995).

In another example, the first flow rate amount may be a first profile of flow rates and the second flow rate amount may be a second profile of flow rates. In other words, the velocities at the first and second measurement positions may be measured periodically, the length of the period may depend on the accuracy required by the application.

Each velocity value of the first and second velocity profiles may then be compared in order to calculate a plurality of time delays across the profiles. The time delays may then be averaged to arrive at the last time delay to be used to calculate the PWV.

In further examples, obtaining the first and second distributions of flow rates may include obtaining first and second frequency spectra, respectively.

In other words, a frequency spectrum is obtained at each measurement depth, which in this case represents z₁And z₂The distribution of the flow velocity. The frequency of the obtained doppler signal will include a frequency offset compared to the originally generated ultrasound pulse, which is directly related to the flow velocity at the measurement point. As the velocity at a given measurement point changes over time, the frequency offset in the resulting doppler signal will change, resulting in a spectrum of different received frequencies. Frequency binning may then be taken to represent a given speed or range of speeds.

In existing systems, the peak flow velocity may be determined by spectral doppler methods, which involve the calculation of the frequency spectrum of the doppler signal over an ensemble of ultrasound pulse transmissions, which in some applications may be up to 256 transmissions. A higher number of emissions increases the speed resolution that can be observed; however, it also reduces the temporal resolution with which the difference in velocity can be observed.

In such cases, the time resolution is not as critical as it is for the proposed PWV measurement method described above, where a significantly lower number of emissions (such as 16) may be beneficial. In fact, these two methods can be used in parallel for simultaneous flow rate and PWV measurements from the same raw measurement data.

In step 150, a time delay is calculated based on the first flow rate metric and the second flow rate quantity.

As discussed above, the time delay may be calculated based on an average first velocity at the first measurement point and an average second velocity at the second measurement point.

Alternatively, the time delay may be calculated by determining an individual time delay for each speed measurement across the distribution of the first and second speed measurements and calculating an average time delay based on this. The individual velocity measurements may take the form of a range of velocities of the velocity profile or a frequency bin of the spectrum of doppler signal frequencies.

As discussed above, changes in velocity over time may include characteristic features that relate to, for example, the heart cycle as it propagates through the vascular system.

As such, the time delay may be calculated based on a comparison between a feature of a first distribution of velocities (referred to as a first feature) and a feature of a second distribution of velocities (referred to as a second feature). The first and second features need not be singular features but may comprise a plurality of different features occurring in both the first and second velocity distributions.

By way of example, the first and second features may include one or more of: instantaneous peak velocity; starting a pulse; a peak time; and a maximum acceleration. In particular, the instantaneous peak velocity represents the maximum flow velocity, which is typically found in the center of the vessel lumen. This feature is less affected by irrelevant parts of the velocity profile, such as may be indicative of wall motion rather than the slow velocity of blood flow.

First and second velocity metrics z₁And z₂The time delay between can be calculated using a cross-correlation method.

The cross-correlation method of measuring the similarity between two signals can be performed directly on the obtained doppler signal before any frequency spectrum calculation. Thus, with this method, no frequency spectrum needs to be calculated. Instead, the time delay representing the velocity of the pulse wave is determined by finding z at which₁And z₂The doppler signal in between provides the time delay of the highest correlation to be obtained.

In step 160, the pulse wave velocity is calculated based on the time delay.

Due to z₁And z₂The distance x between is known and the time delay Δ t is calculated using one of the various methods arranged above, it is a simple matter to determine PWV as:

in addition to the velocity metric, a pressure metric may also be obtained intravascularly for use when calculating the pulse wave velocity as explained above with reference to the following equation:

wherein the content of the first and second substances,

by calculating the PWV via two methods, it is possible to increase the accuracy of the last PWV and to check the quality of the measurement based on the velocity metric.

Fig. 3 shows an intravascular device 200 aligned with a central axis 210 of a blood vessel 220.

The intravascular device generates a plurality of ultrasound pulses 230 and receives a first distance z from the device₁A plurality of echo signals 240 from a first measurement location, and a second distance z from the device₂A plurality of echo signals 250 of the second measurement location. The first and second distances are separated by a distance x.

A plurality of echo signals 240 from a first measurement location are used to form a first doppler signal and a plurality of echo signals 250 from a second measurement location are used to form a second doppler signal.

FIG. 4 shows the flow velocity v_fAnd time graph 260.

The graph shows a first velocity profile 270 and a second velocity profile 280 acquired from the first and second doppler signals, respectively. As can be seen from the figure, the two patterns are of the same shape, the second pattern being offset in time. This shift in time represents the time it takes for the pulse to travel along the vessel and is therefore indicative of PWV. By comparing the first and second patterns, it is possible to derive a time delay Δ t that can then be used in combination with the known distance x to derive PWV.

Figures 5 and 6 each show 4 graphs relating to doppler signals obtained over time T, with the graph of figure 6 showing the 20ms portion of the graph of figure 5 between 7830ms and 7850 ms.

Directing attention to fig. 5, a first graph 300 and a second graph 310 show graphs of the separated I and Q components, respectively, of a received ultrasonic doppler signal. In other words, the first and second graphs show the I and Q components of the doppler signal following phase-quadrature demodulation. Each of the first and second graphs represents a different measurement depth, or region of interest (ROI).

Each time point of the graph results from one pulse transmission caused by the intravascular device. In this example, the pulse repetition rate is 50kHz and is therefore the sampling frequency of these signals and ultimately also the estimated speed signal from which the PWV is calculated. The third graph 320 and the fourth graph 330 show the instantaneous velocity as estimated by the Kasai (third graph) and Loupas (fourth graph) algorithms. In other words, the third graph shows the estimated Kasai velocity V_KAnd the fourth graph shows the estimated Loupas speed V_L. These algorithms estimate the instantaneous average frequency in the I and Q signals that is directly related to the average flow velocity in the vessel by means of the doppler equation. These two graphs indicate that the speed changes from close to 0 to roughly 0.2m/s just before 7850 ms. This change in velocity is due to the pulse wave and looking at the signal carefully, it can be seen that the velocity changes at a low depth (3-4mm) before it changes at a higher depth (e.g. 8-9 mm). In other words, there is an observable time delay between the pulse waves reaching a measurement depth of 3-4mm and a measurement depth of 8-9 mm.

Fig. 6 shows the graph of fig. 5 between times 7830ms and 7850 ms. The fifth graphic 340 corresponds to the first graphic 300. The sixth graphic 350 corresponds to the second graphic 310. The seventh graphic 360 corresponds to the third graphic 320. The eighth graphic 370 corresponds to the fourth graphic 330.

These graphs show the time around the arrival of the pulse wave, from which it can be seen that there is a time delay of a few milliseconds between the arrival of the velocity wave between different depths in the seventh graph 360 and the eighth graph 370. Further, the fifth 340 and sixth 350 graphs show the changes in the I and Q signals resulting from the velocity wave, where the high frequency components appear at the arrival of the velocity wave around 7840 ms.

Fig. 7 shows an alternative graphical representation 380 of the Loupas speed shown in the eighth graph 370 of fig. 6.

The velocity wave is shown by shading as a function of time T (horizontal axis) and distance or depth D (vertical axis) from the transducer. Graphically, pulse wave velocity can be interpreted from this image by plotting lines along the phase of the velocity wave. Is equivalent toThe slope of this line indicates the pulse wave velocity.

Fig. 8 shows a graph 390 of cross-covariance X-co versus time T, which is equivalent to the cross-correlation after subtracting the average from the signal. The cross-covariance is calculated between the velocity signal obtained at the 5.5-6.5mm measurement depth, identifiable by the graph 395 with the peak centered at 0ms, and the velocity signal obtained at the remaining measurement depth. The position of the peak in the cross-covariance indicates the delay between the arrival of the pulse wave between the various measurement depths. As an example, the pulse wave at a depth of 8-9mm is most closely related to the pulse wave at a depth of 5.5-6.5mm when the latter is delayed by about 1 ms. Furthermore, the pulse wave at a depth of 3-4mm is optimally related to the pulse wave at a depth of 5.5-6.5mm when the predecessor is delayed by approximately 1.2 ms.

FIG. 9 shows the optimal lag time T for the maximum cross-covariance between any two measured depths (on the x-axis) plotted against the known distance between the measured depths (on the y-axis)_lagThe graph 400. The data points convey a trend indicative of the pulse wave velocity shown by the dashed line. The pulse wave velocity can be estimated as the slope of this line, which in this example results in a value of 2.1 m/s.

The intravascular device may comprise a plurality of ultrasound transducer elements, in which case the intravascular ultrasound pulses may be directed along a central axis of the blood vessel using electronic beam steering and/or electronic beam focusing. In this way, the measurement results may be made less sensitive to the orientation of the intravascular device.

Electronic beam steering and focusing may be applied to optimally align the ultrasound beam with the flow direction in order to ensure that the axis on which the pulse wave velocity is evaluated corresponds to the axis of pulse wave propagation.

This can be achieved by optimizing the ultrasound beam angle so as to maximize the strength of the doppler signal, for example, at the point where the flow velocity is greatest. The optimization of the beam angle may be performed in an iterative and/or adaptive manner.

Furthermore, the measurements may be performed with two separate ultrasound transducers, which may be implemented on two separate intravascular devices. This may provide a greater degree of freedom in selecting the distance between the two measurement positions.

Other variations to the disclosed embodiments can be understood and effected by those skilled in the art in practicing the claimed invention, from a study of the drawings, the disclosure, and the appended claims. In the claims, the word "comprising" does not exclude other elements or steps, and the word "a" or "an" does not exclude a plurality. The mere fact that certain measures are recited in mutually different dependent claims does not indicate that a combination of these measures cannot be used to advantage. Any reference signs in the claims should not be construed as limiting the scope.