A portable ultrasound device and a method for its operation

Abstract

A portable ultrasound device (100) for non-invasively imaging a selected sub-cutaneous structure, comprises: (a) a plurality of parallel transducer arrays (102) of transducer elements, each array (102) being obliquely angled and each transducer element comprising a transmitter transducer and a receiver transducer for continuously transmitting acoustic waves at ultrasonic frequency in a carrier signal having a predetermined frequency range toward a patient's body and continuously receiving echo signals in a predetermined frequency range from the patient's body following reflection of ultrasound energy, the plurality of arrays (102) of transducer elements enabling imaging of the sub-cutaneous structure (1104, 1105) in multiple transverse and lateral planes; (c) an acoustic lens (103) for manipulating transmission of the acoustic waves; and (d) a transducer controller (104) for operating in continuous wave doppler mode and communicable with a processor (109) for processing received echo signals from the transducer arrays (102). The transducer controller (104) controls a scanning order and scanning rate of transducer elements within the plurality of transducer arrays (102) for generating and processing a Doppler signal enabling imaging of the sub-cutaneous structure (1104, 1105) on the device display (110).

Description

[0001]A PORTABLE ULTRASOUND DEVICE AND A METHOD FOR ITS OPERATION TECHNICAL FIELD The present invention relates to a portable ultrasound device and method of operation. BACKGROUND ART The following discussion of the background art is intended to facilitate an understanding of the present invention only. The discussion is not an acknowledgement or admission that any of the material referred to is or was part of the common general knowledge as at the priority date of the application. Rapid and reliable vascular access using a vascular access device, typically a cannula, is an essential step in the treatment of patients. The overwhelming majority of vascular access devices rely on the clinician’s ability to feel and see the vessel, causing an issue in the reliability of accessing the vessel. Ultrasound (US) guidance improves the reliability of vascular access by letting clinicians visualise the vein with technology. However, US-guided vascular access requires the presence of clinicians or others with advanced training to control and interpret images from traditional ultrasound devices (i.e. B-mode image). Although miniaturization of US systems has advanced accessibility of US to clinicians, most US systems used are still trolley style systems that are not easily accessible. Continuous Wave Doppler (CWD) ultrasound is a candidate for imaging underlying vessels for clinicians and is most used to monitor blood flow in vessels by detecting a Doppler signal, which can help clinicians identify potential issues like vessel occlusion. The Applicant’s WO 2022221913 discloses a portable ultrasound device having a plurality of obliquely angled parallel arrays of transducer elements operable in continuous wave doppler mode to generate an image of a sub-cutaneous structure. The arrays of transducer elements allow imaging of the sub-cutaneous structure in multiple transverse and lateral planes. While the Applicant’s portable ultrasound device represents a significant advance, there is opportunity for enhancement of Doppler signal processing to provide either faster scanning or higher resolution scanning dependent on the requirements of the user. The present invention has been developed against this background. SUMMARY OF INVENTION In a first aspect, the present invention provides a portable ultrasound device for non- invasively imaging a selected sub-cutaneous structure in a subject, comprising: (a) a housing; (b) a plurality of arrays of transducer elements, each array being obliquely angled and arranged in parallel and each transducer element comprising a transmitter transducer and a receiver transducer, located within said housing for continuously transmitting acoustic waves at ultrasonic frequency in a carrier signal having a predetermined frequency range toward a body of a subject and continuously receiving echo signals from the body of the subject following reflection of ultrasound energy, said plurality of arrays of transducer elements enabling imaging of the sub- cutaneous structure in multiple transverse and lateral planes; (c) an acoustic lens disposed relative to said plurality of arrays of transducer elements for manipulating transmission of said acoustic waves; (d) a transducer controller for operating said plurality of arrays of transducer elements in continuous wave doppler mode and communicable with a processor, operable under control of electronic program instructions, for processing said received echo signals from said plurality of arrays of transducer elements, wherein said transducer controller is operable, under control of the electronic program instructions, to control a scanning order and scanning rate of a plurality of transducer elements within said plurality of arrays within said plurality of arrays of transducer elements for generating a Doppler signal processable to remove interference and provide an image of the sub-cutaneous structure. Preferably, the parallel arrays of the ultrasound device are angled at an oblique angle of insonation θ where 45<θ<80° to ensure the Doppler effect is captured. With obliquely angled transducers, an air gap is created between the face of a transducer element and a tissue surface, causing attenuation of the signal and a greatly reduced receive signal. While ultrasound gel has been a traditional response to the air gap, this is not an ideal solution for a plurality of transducers. The acoustic lens of the above ultrasound device enables manipulation of the transmitted acoustic waves to optimise Doppler shift signals for underlying sub-cutaneous structures. The ultrasound device has at least one continuous wave (CW) mixer for mixing the carrier signal with the received echo signals. The processor executes a sequential scan pattern for the transducer elements. The sequential scan pattern may comprise activating a first transducer element for a dwell period sufficient to capture a Doppler signal, deactivating the first transducer element at the end of the dwell period and activating a second transducer element. The sequential scan pattern should be repeated until all of the transducer elements in an array have been activated for the dwell period. A plurality of transducer elements, optionally a pair of transducer elements, may be activated in parallel with the receive signals being input to a continuous wave mixer with Doppler signals from the plurality of transducer elements. More receive transducer elements may be activated than transmit transducer elements during a scan. The transducer controller may be used to control scanning rate, optionally by controlling dwell period of activation of a transducer element during scanning as described herein. Optionally, a user interface may be included for a user to control scanning rate. The ultrasound device may have a plurality of CW mixers with the advantage that scan rate is increased over the case where the ultrasound device has a single CW mixer because a plurality of transducer elements can be activated concurrently. The reduction in scan time is apparent from the formula: Scan time = No. elements / No. CW mixers*dwell period With a plurality of CW mixers, the controller concurrently activates each of a selected plurality of transducer elements in a scan pattern executed by the controller. One CW mixer may be provided per transducer array (i.e. 1:1 ratio of CW mixer to transducer array) though the ratio may be greater than 1:1 dependent on space constraints on-board the portable ultrasound device In one embodiment, the number of CW arrays is the same as the number of transducer arrays. Again, the controller activates the selected plurality of transducer elements for a dwell period sufficient to capture a Doppler signal. Preferably, each received echo signal is combined with the carrier signal by a dedicated CW mixer. The scan patterns executed by the transducer controller are conducted to minimise cross- talk or acoustic interference between different, particularly adjacent transducer elements. To that end, the activated transducer elements are separated by at least one inactive transducer element during a scan. However, in an alternative embodiment, where the ultrasound device has one CW mixer, which has an advantage of minimising electronic area, the transducer controller may activate a plurality of physically adjacent transducer elements. While this may reduce resolution, an advantage is obtainable in terms of a faster scan rate. Due to the CW mixer summing all transducer element input signals, a group of activated transducer elements acts as a single larger transducer element resulting in lower resolution. On the other hand, less transducer elements are sequentially scanned, decreasing the time to complete a scan. Further reduction in scan time is achievable by providing a plurality of CW mixers and activating transducer elements in groups either automatically or on user selection through a user interface. In such embodiments, the controller may execute a scan comprising concurrent activation of selected groups of transducer elements of the parallel arrays. The selected groups of transducer elements may comprise physically adjacent transducer elements conveniently arranged in a linear transducer array. A first group of transducer elements may be activated in a first sequential scan pattern and, on detection of a selected sub-cutaneous structure, by processing of the received echo signals from a second group of transducer elements comprised within said first group of transducer elements, activates transducer elements within said second group of transducer elements in a second sequential scan pattern; and wherein the second group of transducer elements has a lesser number of transducer elements than the first group of transducer elements. The ultrasound device may be configured to enable switching between the first group of transducer elements and the second group of transducer elements, for example by including a switch for each of the groups of transducer elements to allow control over field of view by changing the number of activated transducer elements. The transducer controller may enable an adaptive scan pattern utilising a plurality of scan protocols, each scan protocol containing an activation configuration and a corresponding flow reconstruction algorithm in the case of blood vessels. Optionally, the transducer controller may use a first scan protocol and a second scan protocol, the first scan protocol having an activation configuration with a short dwell period optimised for scan rate, optionally less than 50 ms, optionally less than 30 ms, optionally greater than or equal to 10 ms, optionally 20 ms. The second scan protocol may have an activation configuration with a relatively longer dwell period optimised for displaying a blood vessel, optionally with flow direction information. The longer dwell period may range from about 20 to greater than 200ms with the minimum longer dwell period being the minimum scan time to obtain flow direction in vasculature with low blood flow in a slow flow vein for example. The ultrasound device may be configured to enable a user to switch between the plurality of scan protocols, optionally through a user interface. In preferred embodiments, the user may change field of view by command to the transducer controller by a user interface, optionally the display where a touch screen or by other input means. For example, the transducer controller may be commanded to activate a first group of transducer elements in a first sequential scan pattern (optionally in one dimension) and activate a second group of transducer elements in a second sequential scan pattern (optionally in a second dimension). The user may also switch between different viewing options, for example one dimensional and multi-dimensional, conveniently two dimensional. For example, the user may choose to activate a first group of transducer elements, for example one transducer array detecting the sub-cutaneous structure. While this may limit the vascular information captured to one dimension, it reduces the number of active transducer elements required to complete a scan, reducing the scan completion time. The user may then switch to activate all transducer arrays (a second group of transducer elements with a greater number of transducer elements than the first group of transducer elements in this instance)) to visualise the sub-cutaneous structure in two-dimensions at a slower scan rate. In the above embodiments, the processor may compare the received echo signal from the first sequential scan pattern with the received echo signal from the second sequential scan pattern to determine if the ultrasound device has moved. On determination that the ultrasound device has moved, the controller activates a third group of transducer elements, the third group of transducer elements containing a greater number of transducer elements than the second group of transducer elements. If the ultrasound device is determined not to have moved, the controller activates a third group of transducer elements, the third group of transducer elements containing a lesser number of transducer elements than the second group of transducer elements. The processor is able to follow any desired demodulation process. Conveniently, quadrature demodulation forms part of the processing. In this case, the received echo signals from the receiver transducer elements are mixed with the carrier signal by the CW mixer(s) to remove the carrier signal by quadrature demodulation. A mixer output signal has the following result or equation for the I-channel and Q-channel after removal of the high frequency components: Where VDC(t) is the DC offset, Amis the mixer sinusoid amplitude, Adis the Doppler sinusoid amplitude, ^^is the Doppler frequency, ^^is the Doppler phase shift and ^^is the mixer phase shift. The mixer output signal has a DC offset, substantially larger than the useful Doppler signal representative of blood flow and vasculature, that must be removed or compensated for to isolate the Doppler signal for further processing. The DC offset is typically variable with each transducer element activation. The DC offset could be removed by a high pass filter, though this would undesirably compromise scanning rate, or most preferably by DC offset removal modality which avoids requirement for a high pass filter. More broadly, the processor preferably isolates the Doppler signal from the mixer output signal by subtracting the DC offset signal from the mixer output signal. In one embodiment, the DC offset removal circuitry comprises: a first sample and hold circuit for measuring the DC offset signal, hold the DC offset and redrive the hold signal; and a difference amplifier that subtracts the measured DC offset signal from the mixer output signal to isolate the Doppler signal. Alternatively, the processor may sample the output signal for each activation of transducer element(s) to separately determine a DC offset signal for each activation, determine the DC offset signal for each sample and subtract the DC offset signal from the output signal to isolate the Doppler signal. The DC offset signal may also drift as a function of fluctuation in received echo signals. The processor may be configured to determine drift in DC offset signal as a function of fluctuation in received echo signals, the drift being subtracted from the Doppler signal to further isolate the Doppler signal. The processor advantageously processes one or more components of the Doppler signal representative of blood flow to provide a representation of the sub-cutaneous structure, in particular vasculature. One or more components may be selected from the group consisting of digital raw voltage Doppler signal, a magnitude-frequency spectrum of the Doppler signal and the phase-frequency spectrum of the Doppler signal. For example, the processor may process the magnitude-frequency spectrum. Preferably, the processor processes both the magnitude-frequency spectrum and the phase-frequency spectrum of the Doppler signal. Whatever the selected components, the objective is to obtain values that are representative of the sub-cutaneous structure, in particular vasculature. Conveniently, the processor processes the component with a Fast Fourier Transform (FFT) analysis. A magnitude-frequency spectrum may be processed by a magnitude- FFT analysis. A phase-frequency spectrum may be processed by a phase-FFT analysis. Additionally, or alternatively, the processor may also process received ultrasound data in the time domain. Before component processing, for example by FFT analysis as above described, signal interference should be removed. Interference may be induced by soft tissue movement or by other factors. To this end, a smoothing method is desirably conducted prior to FFT analysis. Following removal of interference, the processor may execute an extraction algorithm to filter the Doppler signal to extract signals in the frequency and / or time domain. For example, the extraction algorithm may extract signals at selected frequency representative of a sub-cutaneous structure, in particular vasculature. Following execution of the extraction algorithm, the processor executes a reconstruction algorithm for providing a representation image of the sub-cutaneous structure. Such reconstruction algorithm may use processed frequency domain and / or time domain data as an input. Vessel characteristics may also be an input, e.g. blood vessel diameter, blood vessel diameter, blood vessel depth and blood vessel ellipticity though not limited to these. For example, the representation of the sub-cutaneous structure may be in the form of an array, a line or a heat or colour map linking the processed signals for transducer elements that have either detected the sub-cutaneous structure or have a high probability of having detected the sub-cutaneous structure. The image is desirably displayed, preferably as a coronal view, on a screen of the portable ultrasound device. The display may provide one dimensional data representative of the sub-cutaneous structure. In a further aspect, the present invention provides a method for imaging a sub-cutaneous structure in a subject with the above described portable ultrasound device. In a still further aspect, the present invention provides a method for imaging a sub- cutaneous structure in a subject with a portable ultrasound device, said method comprising: (a) continuously transmitting acoustic waves at ultrasonic frequency in a carrier signal having a predetermined frequency range toward a body of a subject from a plurality of arrays of transducer elements, each array being obliquely angled and arranged in parallel and each transducer element comprising a transmitter transducer and a receiver transducer; (b) manipulating transmission of said acoustic waves; (c) continuously receiving echo signals from the body of the subject following reflection of ultrasound energy; (d) operating said plurality of arrays of transducer elements in continuous wave doppler mode; and (e) processing said received echo signals from said plurality of arrays of transducer elements, wherein a scanning order and scanning rate of a plurality of transducer elements within said plurality of arrays of transducer elements is controlled for generating a Doppler signal processable to remove interference and provide an image of the sub-cutaneous structure. Advantages of the invention include providing a readily interpretable image of a sub- cutaneous structure, in particular vasculature to assist effective cannulation. The image may be generated from a Doppler signal that is strong, stable and free from interference in a fast manner. In this regard, the cannulation procedure – for example – typically requires to be completed and efficiently, in particular urgently. BRIEF DESCRIPTION OF THE DRAWINGS Further features of the portable ultrasound device and method for imaging of the present invention are more fully described in the following description of several non-limiting embodiments thereof. This description is included solely for the purposes of exemplifying the present invention. It should not be understood as a restriction on the broad summary, disclosure or description of the invention as set out above. The description will be made with reference to the accompanying drawings in which: Figure 1 is a block diagram of a portable ultrasound device according to one embodiment of the present invention. Figure 2 is a schematic representation of a transducer scan pattern where the portable ultrasound device of Figure 1 has a single continuous wave mixer. Figure 3A is a schematic representation of a transducer scan pattern where the portable ultrasound device of Figure 1 has a plurality of transducers and a plurality of continuous wave mixers. Figure 3B is a schematic representation of a transducer scan pattern where the portable ultrasound device of Figure 1 has a single transducer and a plurality of continuous wave mixers. Figure 4 is a schematic representation of a transducer scan pattern for tuning of resolution in the portable ultrasound device of Figure 1 having a single continuous wave mixer. Figure 5 is a flow chart for a transducer scan pattern with a variable resolution and scan time. Figure 6A is a schematic diagram of a first embodiment of acoustic lens for use in the portable ultrasound device of Figure 1. Figure 6B is a schematic diagram of a second embodiment of acoustic lens for use in the portable ultrasound device of Figure 1 that allows focusing for a pre-angled lens. Figure 6C is a schematic diagram of a third embodiment of acoustic lens for use in the portable ultrasound device of Figure 1. Figure 7A is a high-level block diagram of an DC offset correction unit forming part of the processor of the portable ultrasound device of Figure 1. Figure 7B is a detailed diagram of the DC offset correction unit implementation. Figure 8 is an example of a Doppler frequency shift profile for the portable ultrasound device including the DC offset correction unit of Figure 1 and representing paired values in the magnitude-frequency (Figure 8A) and phase-frequency (Figure 8B) domains. Figure 9A is a flowchart of a magnitude Fast Fourier Transform (FFT) analysis of the Doppler frequency shift profile of Figure 8. Figure 9B is a flowchart of a phase FFT analysis of the Doppler frequency shift profile of Figure 8. Figure 10A is a flow diagram showing one embodiment for processing the Doppler signal from the portable ultrasound device of Figure 1. Figure 10B shows the output of the processing of Figure 10A providing a representation of a sub-cutaneous structure on a display of the portable ultrasound device of Figure 1. Figure 11A is a flow diagram of a further embodiment for processing the Doppler signal from the portable ultrasound device of Figure 1. Figure 11B shows the output of the processing of Figure 11A providing a representation as a line of a sub-cutaneous structure on a display of the portable ultrasound device of Figure 1. Figure 12A is a flow diagram of a still further embodiment for processing the Doppler signal from the portable ultrasound device of Figure 1. Figure 12B shows the output of the processing of Figure 11A providing a representation as a colour map of a sub-cutaneous structure on a display of the portable ultrasound device of Figure 1. Figure 13 is a flowchart of a scan pattern utilising multiple scan protocols. Figure 14 is a graphical example of the Doppler signal post CW mixing and the response of a high pass filter to a DC step. DESCRIPTION OF PREFERRED EMBODIMENTS Referring to Fig.1, there is shown portable ultrasound device 100 configured to display sub-cutaneous structures on display 110 disposed on the housing. Portable ultrasound device may take the form as described in the Applicant’s Australian Patent No. 2022261790 or the Applicant’s co-pending PCT Application No. PCT / AU2025 / 051046 , the contents of which are incorporated herein by reference. In some embodiments, ultrasound device 100 can be used in the medical field for inspection of sub-dermal structures, including but not limited to, vascular vessels such as arteries and veins. In one embodiment, such an inspection is useful to healthcare professionals (HCP) inserting a cannula into a vascular vessel in a cannulation procedure. This embodiment aids the HCP in finding a suitable vasculature and in aligning the cannula. Those skilled in the art of vascular access devices will understand that a cannula can be replaced with a needle for a venipuncture procedure. Further description of preferred embodiments is directed to imaging of vasculature, such as veins for cannulation. Portable ultrasound device 100 operates in continuous wave doppler (CWD) mode which enables a high sensitivity to blood flow. Though the present embodiment envisages CWD mode as herein described as the sole mode of visualisation, it will be understood that ultrasound device 100 may, in other embodiments, be configured to enable other modes of operation, such as A mode, B mode, C mode, M mode, pulsed wave doppler mode and colour doppler mode. Ultrasound device 100 comprises a plurality of parallel transducer arrays 101, each array 101 being linear and comprising a plurality of transducer elements 102. The transducer elements 102 may be as known in the art of acoustics, for example taking the form of piezo transducers, cMUT transducers and pMUT transducers but not limited to these. Three transducer arrays may be provided in one embodiment, this enabling efficient imaging and visualisation, though the drawings may show one or two transducer arrays for purposes of illustration. Another number of transducer arrays may be selected. The plurality of parallel transducer arrays 101 provides the ability to image vasculature in multiple transverse planes to display the lateral position of the vessel on a screen 110 (see Figs.10-12) of ultrasound device 100, for example as it is moved across the skin, for example on the forearm, of a patient. The parallel transducer arrays are obliquely angled at an angle of insonation, θ, where 45<θ<80° to ensure the Doppler effect is captured. The parallel transducer arrays 101 are spaced apart a distance ξ enabling desired visualisation or imaging (these terms being used interchangeably) of a vein to determine its direction and linearity over a distance. This distance ξ must be optimised to reduce bulk of the ultrasound device 100 and a distance of 5-30 mm between transducer arrays 101 is preferred based on currently achievable resolution. Shorter distances may be achievable. The transducer array 101 transfers data back and forth with transducer controller 104 for processing by processor 109. Transducer controller 104 controls the scanning order and scanning rate of selected transducer elements 102. The transducer controller 104 executes, being programmed with suitable electronic instructions to do so, a predetermined scan pattern or algorithm as described further below. The scan algorithm allows improvement in signal strength, stability, interference characteristics and scan rate. Ultrasound device 100 further comprises a DC offset correction unit (DOCU) 105 configured to receive processed analog signals from a continuous wave (CW) mixer chip and remove the large DC offset signal in very short time (near 0) compared to a high pass filter (around 100 ms). In some embodiments, ultrasound device 100 comprises one CW mixer chip and in other embodiments, ultrasound device 100 comprises a plurality of CW mixer chips. Commonly used CW mixers, such as analog front-end (AFE) chips, voltage- controlled amplifier chips and voltage-gain amplifier chips, may be used in ultrasound device 100 in any of the embodiments described. DOCU 105 has the function of isolating and amplifying the Doppler signal present in the processed analog signal and converting the analog Doppler signal into digital data. In one embodiment, the digital data is further transformed to a frequency domain representation of the Doppler signal using a Fast Fourier Transform (FFT) analysis of component(s) of the Doppler signal, resulting in an output signal having a magnitude- frequency component and a phase-frequency component. Processor 109 of ultrasound device 100 is configured to process received CWD signals. In particular, processor 109 processes components of the CWD signal to enhance detection and imaging of underlying vasculature. Preferably, the processor 109 processes a digital raw voltage CWD signal. More preferably, the processor 109 processes at least one of, and preferably both of, the magnitude-frequency spectrum of the CWD signal and the phase-frequency spectrum of the CWD signal. Conveniently, the selected components of the CWD signal are analysed by Fast Fourier Transform (FFT) analysis forming part of an extraction algorithm executed by processor 109. The magnitude-FFT analysis method is denoted 106 and the phase-FFT analysis is denoted 107, where the main objective of the FFT analysis is extracting values from each component that is representative of vasculature. In further embodiments, the extracted values are not only representative of the presence of vasculature but also characteristics of the vasculature which are not assumed constant during scanning. Such characteristics may, without limitation, be selected from the group consisting of blood vessel velocity, blood vessel direction, blood vessel depth and blood vessel diameter. The extracted values are inputted to a reconstruction algorithm 108 that has the function of reconstructing vasculature and displaying the data in an easily interpretable manner for end users on display 110. The reconstruction algorithm 108 allows visualisation of vasculature, preferably as a coronal view of the vasculature, such as a vein for cannulation. Alternatively or additionally, the direction of the vein may be visualised or imaged and displayed on display 110. Fig.2 illustrates a transducer scan pattern as controlled by the transducer controller 104 where only one CW mixer chip is utilised. In this exemplary embodiment, there are two separate linear transducer arrays, each comprising a plurality of transmit and receive element pairs 201 (such pair 201 being referred to, for convenience, as a transducer element). However, it is to be understood that transducer controller 104 may implement scan patterns for any number of transducers or transducer elements. While linear transducer arrays are convenient for imaging as described herein, it is to be understood that transducer elements or arrays may be arranged with differing geometry in alternative embodiments. In the embodiment of Fig.2, noting that alternative scan patterns are described below, the scan begins with a transducer element 201 within transducer array 202 being activated to continuously transmit and receive acoustic waves (i.e. continuous wave doppler) for a set period, known as a dwell period controllable, in all embodiments, by the transducer controller 104. The dwell period is required to ensure a period of the Doppler shift frequency of interest is acceptably sampled. The dwell period is preferably less than 100 ms, more preferably less than 50 ms, or most preferably less than 30 ms, for example 20 ms enabling collection of signals representative of a slow flow vein. A lower dwell period maximises scan rate (subject to transducer controller 104 capacity and processor 109 processing capacity) as each transducer element is activated for a less time while still being able to capture the Doppler shift frequency, for example for slow flow veins. However, such fast scan rates create a challenge in removal of DC offset which is described together with embodiments providing DC offset removal in the context of such fast scan rates. Once the transducer element 201 has transmitted and received the acoustic waves for the dwell period, it is deactivated and an adjacent transducer element is activated, again transmitting and receiving the acoustic waves for the determined dwell period as described above. The adjacent transducer element is spaced a short distance, e.g.1.2 mm distance from centre of transducer element 201 to adjacent transducer element, which enables efficient resolution for visualising or imaging the pathway of a sub- cutaneous structure, in particular vasculature. This transducer element spacing may be selected as desired to achieve this objective. This activation process is repeated in all adjacent transducer elements, creating a directional scan 203 within the transducer arrays as indicated by the arrows in Fig. 2. Once all elements within a transducer array have been activated, controller 104 causes the same scan pattern to recommence in an adjacent transducer array 204 spaced a short distance, between 5 and 30mm from transducer array 202. After all transducers and elements have been activated, the scan pattern is repeated. Advantageously, controller 104 commands a delay between deactivation and activation of adjacent transducer elements to reduce the occurrence of acoustic cross interference between adjacent transducer elements, where the transmit of one transducer element is captured by an adjacent transducer element receive. Acoustic cross interference results in an increase of interference due to undesired acoustic energy being received by the wrong transducer element. Preferably, the delay would be substantially less than 100 ms, more preferably less than 50 ms or most preferably less than 10 ms. In another embodiment, the activation of transducer elements 201 is completed in a scatter pattern where transducer elements are not activated in a sequential manner. Instead, the next transducer element activated is not physically adjacent to the previous activated transducer element and this process is completed until the scatter pattern scan, involving activation of all transducer elements is complete. The scatter pattern has the advantage of spatially separating sequentially activated transducer elements to minimise the occurrence of acoustic cross interference. Figs.3A and 3B illustrate a scan pattern controlled by the transducer controller 104 where the portable ultrasound device 100 utilises a plurality of CW mixer chips. In the embodiment of Fig.3A, there are shown two separate parallel linear transducer arrays 102 comprising transducer elements 201 and two CW mixer chips. However, subject to processing capability of transducer controller 104 and processor 109, the scan pattern may be applied for any number of CW mixer chips. The scan pattern of Fig. 3A involves parallel activation of transducer elements by transducer controller 104 where an active transducer element 301 on the first transducer 102 outputs its receive signal to the first CW mixer chip and an active transducer element on the second transducer 303 outputs its receive signal to the second CW mixer chip. Similarly to the scan pattern in Fig.2, after activation of the transducer elements for the determined dwell period (as described above) is complete, the transducer elements 301, 303 are deactivated and the physically adjacent transducer elements 306, 307 are activated. This parallel activation of transducer elements occurs simultaneously, advantageously reducing the time to complete a full scan. The scan pattern of Fig.3B involves an arrangement where a plurality of CW mixer chips are utilised on a single transducer. In this embodiment, transducer controller 104 commences a scan by activating, in parallel, a transducer element connected to the first CW mixer chip 301 and a transducer element connected to the second CW mixer chip 303. Transducer controller 104 then activates a transducer element 302 or 304 that has not been activated in the scan. An important consideration is avoidance of parallel activation of transducer elements thus transducer elements are spaced apart by a spacing 305 to prevent acoustic cross interference between activated transducer elements. Preferably, spacing 305 is greater than 5 transducer elements, more preferably larger than 3 transducer element and most preferably equal to or larger than 1 transducer element in an x dimension. This spacing 305 may also be applied to an array of transducer elements where there is minimal spacing between transducer elements in the x and y dimension. Additional CW mixer chips may be used to increase the number of elements activated at any given time within a single transducer and adjacent transducers. It will be understood that the number of CW mixer chips is constrained by the complexity of the electronic transmit pathways, electronic size constraints and the introduction of Acoustic cross interference between transducer elements. Fig. 4 illustrates a scan pattern controlled by the transducer controller 104 where two adjacent transducer elements are activated in parallel and the receive signals are output to a single CW mixer chip. In this scan pattern, transducer elements are activated in groups, in this embodiment and for ease of illustration in pairs. A transducer element 403 is activated at the same time as an adjacent transducer element 402 and two sets of acoustic waves 404, 405 are outputted towards an underlying vessel 401. Received signals from any underlying tissue structure 406 are then received by both transducer elements 402, 403 and processed by the CW mixer chip where the doppler signals are summed. This means that it is not possible to separate which signal is from which spatial location, effectively lowering resolution. However, this has the advantage of multiplying the strength of Doppler shift signals detected facilitating processing to image vasculature. Once the activation of the transducer element pair 402, 403 has finished, activation then occurs for another pair of transducer elements that have not yet been activated in the scan. Advantageously, this scan pattern reduces the time to complete a full scan, with the trade off being on resolution. It will be understood that an activated group of transducer elements may include more than a pair of transducer elements. Additional CW mixer chips may also be utilised to further reduce the time to complete a full scan. Alternatively, when activating multiple parallel transducer elements, more receive elements may be activated than transmit elements. In this scan pattern, transducer element 403 and an adjacent transducer element 402 are activated, however only a single acoustic wave is output 403 towards an underlying vessel 401. Received signals from any underlying tissue structure 406 are then received by both transducer elements 402, 403 and processed by the CW mixer chip where the doppler signals are summed. Advantageously, this scan pattern will increase the strength of the Doppler shift signal detected while not introducing the complexity of timing the two transmission waves to ensure minimisation of destructive interference between the two waves. In another embodiment, the transducer controller 104 includes switches to the elements, to allow for switching between the number of activated adjacent elements at once. This advantageously allows for switching between a reduced time for a full scan with more activated adjacent transducer elements; or a higher resolution with lesser activated adjacent transducer elements. In yet another embodiment, the transducer controller 104 includes switches for the transducer elements, to allow reducing or increasing the field-of-view in both the x and y axis by changing the number of active transducer elements. For example, the switches may only activate one transducer array 102, while this limits the vascular information captured to one dimension, it reduces the active elements required to complete a scan, reducing the scan completion time. In another aspect, the transducer controller 104 contains a user interface to allow the user to control the desired scan pattern. For example, the user may activate only one transducer array 102 to maximise scan rate while visualising or imaging a blood vessel in one dimension and then switch to activate all transducer arrays 102 to visualise or image the blood vessel in two-dimensions at a slower scan rate. Fig.5 illustrates a flow chart outlining the steps completed by the transducer controller 104 for an adaptive scan pattern 501 that varies scan completion time and resolution. The first step is a coarse scan pattern 502 where every transducer element is activated at the lowest resolution (lowest scan completion time) through grouping of transducer elements as described above. Any practical number of transducer elements may be included in a group and any previously presented scan patterns may be adopted in this adaptive scan pattern. After completion of the coarse resolution scan pattern 502, a mask for it is created by processor 109. The mask captures which groups of transducer elements had vascular flow captured from the CWD signal. From the flow mask of elements, a finer scan pattern 503 is completed of transducer elements that had flow captured from the CWD signal. The finer scan pattern 503 involves grouping of a lesser number of transducer elements per group than for coarse scan pattern 502. By grouping a lesser number of transducer elements per group, resolution is higher than during the coarse pattern scan 502. A mask is likewise created for the fine scan pattern 503. Processor 109 then executes a comparison algorithm 504 for comparing the fine scan pattern 503 with the coarse scan pattern 502. The purpose of comparison algorithm 504 is determination of whether ultrasound device 100 has moved such that the previously detected vessel is no longer under the activated transducer elements. Alternatively an accelerometer may be included in the portable ultrasound device to determine device movement. In the embodiment, the comparison algorithm 504 involves an area comparison of flow and no flow between the two masks (coarse and fine). If the flow scan mask of the finer scan pattern is within the flow scan mask of the coarse scan pattern, then the comparison algorithm 504 determines that the ultrasound device 100 has not moved. In another embodiment, the comparison algorithm 504 compares the fine scan pattern flow mask with the number (nt) of activated transducer elements showing flow. If nt represents a majority of transducer elements, then comparison algorithm 504 determines that the ultrasound device 100 has not moved. If the ultrasound device 100 has not moved and the blood vessel remains underneath the same transducer elements, as determined by comparison algorithm 504, then the transducer controller 104 will complete a finer scan pattern on the group of transducer elements detecting flow from the previous scan. The described coarse-fine scanning loop 505 is then repeated until the finest resolution is scanned, one element in an element group. The scanning loop 505 consists of first completing a finer scan pattern on the previous scan pattern’s group of transducer elements capturing flow. Second, the comparison algorithm 504 determines if the ultrasound device 100 has moved as described above. A determination is then made to decrease the resolution if the blood vessel has moved relative to ultrasound device 100 or increase the resolution if the blood vessel has not moved. Once the finest scan pattern 505 has been completed, a comparison is made by comparison algorithm 504 and if the blood vessel has not moved relative to the ultrasound device 100 then the finest scan pattern is repeated. In another embodiment, the transducer controller 104 includes the configuration to vary the dwell time of transducer elements. The transducer controller 104 can reduce the dwell to increase scan rate with the tradeoff of less sensitivity or information about low-speed vessels or alternatively increase the dwell time for the opposite effect. The configurations of both dwell time and scan order are stored in protocols that the transducer controller 104 utilises. In addition, a flow reconstruction algorithm is stored in the protocol that maximises flow reconstruction for the corresponding dwell and scan pattern configuration. Fig.13 illustrates a flow chart outlining the steps completed by the transducer controller 104 for an adaptive scan pattern 1301 that utilises a plurality of scan protocols. As previously described, protocol 1302 contains a transducer activation configuration 1304 and the corresponding flow reconstruction 1305. This is repeated until an input 1308 is given to the transducer controller to change to protocol Y 1303, containing an activation configuration and the corresponding flow reconstruction. It is understood that the adaptive scan pattern 1301 may switch between more than two protocols. In one embodiment, the scan pattern starts in the first protocol 1304 that has an activation configuration 1304 with a short dwell time, preferably less than 20ms, which may be less than the period of the slowest desired flow Doppler shift. The corresponding reconstruction algorithm utilises any techniques described in subsequent descriptions to successfully reconstruct flow from a short dwell period. In this example, a time-based voltage peak-peak approach is utilised to reconstruct the vessel, including blood vessels with flow where a full period is not captured, however no flow direction information is displayed. When an input 1308 is received, the transducer controller 104 then switches to a second protocol 1303. The second protocol 1303 contains an activation configuration 1306 with a longer dwell time, preferably more than 20ms, which contains the period of the slowest desired blood flow for a candidate vein for cannulation. The corresponding reconstruction algorithm utilises any techniques described below to reconstruct flow and acquire directionality information. In this example, a frequency-based approach is used to reconstruct and display the vessel with flow direction information. The input 1308 may originate from a user control input, components within the portable ultrasound device 100 or from an algorithm. For example, the input 1308 may be determined from an algorithm that switches to the longer dwell period protocol when blood vessels are detected. An additional protocol may be included with an activation configuration that is much larger than the preferred dwell period (20ms in the presently described embodiments). The reconstruction algorithm may utilises techniques that are ideal for long dwell period, such as a Short-time Fourier transform, enabling extraction of additional blood vessel information that was not accessible with a short dwell time. The switching 1308 between transducer element activation protocols may be controlled by the user through a user input interface allowing user control or selection between scan rate and information content. This user-controlled flexibility allows real-time adaptation to varying scenarios, such as using the high scan rate protocol while scanning the whole forearm for a suitable blood vessel, for example for cannulation, and then switching to the longer dwell period protocol to collect more detailed vessel-related information when a suitable vein for cannulation is identified. Blood velocity and linearity of the vein are important considerations here. In yet another embodiment, both previously described scan protocols include varying resolution and field of view (FOV) to further enhance scan rate. Where the first protocol 1302 may combine a larger resolution, by activating multiple adjacent transmitter transducer elements, with a short dwell period to increase scan rate. In this embodiment, the switching between scan protocols may be controlled by the transducer controller 104 through an algorithm similar to the one described in Fig 5. The transducer element groups detecting blood flow from the first protocol 1302 are used as the element activation mask for the second protocol 1303 at the finer resolution and longer dwell time. This increases the scan rate of the second protocol 1303 while preserving the more detailed vessel- related information. Figs.6A, 6B and 6C illustrate several embodiments of the transducer array 101 which comprises a plurality of acoustic transducer elements 102 (each including a transmitter and receiver activatable for a dwell period) coupled with an acoustic lens 103. The purpose of acoustic lens 103 is to manipulate transmitted acoustic waves to optimise received Doppler shift signals for underlying vessels. Fig. 6A shows an acoustic lens 103 which reduces the presence of air between the plurality of transducer elements 102 and underlying tissue 603. In this embodiment, three transducer elements 102 are included. When air is present between a transducer element and underlying tissue, attenuation and reflection of the transmitted acoustic waves occur before reaching the underlying blood vessel causing a diminished receive signal. This is further exacerbated for CWD transducers where an angle of insonation is required, causing a gap between the face of the transducers and the underlying tissue. In some embodiments, this gap is replaced by the acoustic lens 103, allowing for the acoustic transducer elements 102 to be at a set angle of insonation 602 (as described in detail above and adopted here) and transmit acoustic waves to the underlying tissue 603 and underlying vessel 604 with minimal attenuation and reflection. Furthermore, the acoustic lens 103 is flat to reduce air gaps and maximise tissue contact. The acoustic lens 103 is manufactured from any material that has a similar acoustic impedance to human tissue such as, but not limited to, silicon or doped silicon. Acoustic lens 103 may comprise a plurality of layers. The angle of insonation 602 and the minimum length of the acoustic wave travel path in the lens 103, known as the ‘lens thickness’, is inversely proportional such that a higher angle of insonation 602 results in a smaller lens thickness. The transducers 102 are set into the acoustic lens 103 at an angle between 45°<θ<80° to minimise the attenuation from the lens 103 while ensuring Doppler signals are still captured. Traditional CWD devices do not allow for control of the depth of scanning. However, in one embodiment, the lens thickness 601 is purposely lengthened to increase attenuation of the acoustic waves to control the maximum penetration depth. This is especially advantageous when visualising underlying veins, more specifically peripheral veins, where the maximum depth would typically be less than 20mm. Alternatively, control of the maximum penetration depth of acoustic waves may be controlled by varying the acoustic impedance of the lens 103. For example, the penetration depth of acoustic waves may be reduced by including dopants to the lens material, varying silicone catalyst to polymer ratio or selecting alternative materials to mismatch the acoustic impedance of the underlying tissue 603 and lens 103. Additionally, varying the acoustic impedance of the lens 103 causes a change in the refraction angle of the transmitted acoustic wave, Refraction angle may be harnessed such that a smaller angle of insonation occurs, resulting in a larger Doppler shift. This has the benefit of maximising signal strength as a larger Doppler shift occurs and a higher transducer angle can be utilised resulting in less loss of the signal strength due to attenuation of the lens. In another embodiment of acoustic lens 103, the surface of the acoustic lens 103 is not flat and is instead shaped to further maximise the contact with underlying tissue 603. An example would be curvature of the surface of the lens 103 to match the contour of the ventral surface of the forearm. Alternatively, the lens 103 mechanical properties could be varied to increase contact surface. For example, if the shore hardness of the lens material is decreased, the lens 103 will be able to conform more easily to the underlying tissue without affecting underlying vascular structures 604. In a still further embodiment of acoustic lens 103, a flat contact surface is not required for locations where the underlying tissue is conformable to the shape. This may occur in regions where the underlying tissue 603 mainly comprises adipose tissue. Referring to Fig. 6B, the lens 103 is the same shape as the transducers 102, reducing the lens thickness 601. Additionally, a curvature is included on the face of the lens 103 to focus the acoustic wave reducing the spread and increase the signal strength. In this embodiment, air is minimised between the transducers 102 and underlying tissue 603 by compressing the lens undulations into the conforming underlying tissue. In a still further embodiment of the acoustic lens 103, layers of materials with differing acoustic properties are utilised. Referring to Fig.6C, the lens 103 comprises two layers of material, where the transducer contacting lens 605 is curved to allow focusing of the acoustic wave and the tissue contacting lens 606 is flat to minimise air between the transducers and the underlying tissue 603. As described above, and in all embodiments, the output signal from CW mixer(s) contains a DC offset signal much larger than the Doppler signal. The DC offset signal must be removed to enable meaningful processing of the Doppler signal. A DC offset correction unit is included for this purpose. Fig.7A illustrates the flowchart of the DC offset correction unit 105. The main function of DOCU 105 is to convert outputs (both in-phase and quadrature) from commonly used CW mixers, such as analog front-end (AFE) chips, voltage-controlled amplifier chips and voltage-gain amplifier chips to a Doppler signal that can be consistently sampled in a fast manner to allow imaging and, through such imaging, make a cannulation procedure, for example, more efficient. Fig.7B demonstrates a detailed diagram of the DOCU 105 implementation alongside the CW mixer output processing circuit 701. The transducer receive element 707 receives a signal, which is passed through a low noise amplifier (LNA) 708. The LNA generates differential signals which are passed to differential in-phase and quadrature mixers 709 and 710. The CW mixers 709, 710 are fed a local oscillator (LO) signal 712, with the quadrature LO signal shifted 90 degrees to the in-phase LO. The output of the mixers 709, 710 are low-pass filtered (LPF) at 711, where the outputs are fed to the DOCU 105. The necessity of DOCU 105 is understood by constructing a simple model of the receive signal and viewing how it is manipulated via the mixing performed in 701. CW mixing is typically performed by passive switch-based mixing with higher order harmonic suppression. This process is mathematically equivalent to multiplying the received signal by a square wave with a frequency equal to the frequency of the transmit signal. The square wave can be expressed by the sum of odd harmonics in a Fourier series, with the first harmonic being a sinusoid at the transmit frequency. The higher order harmonic suppression of the CW mixer filters out the receive signal interaction with the higher harmonics of the square wave, such that the entire mixing process can be modelled as multiplication of the receive signal, ^^^^^^with the fundamental sinusoidal signal at the transmit frequency ^^. Each receive signal is distributed to a pair of CW mixers, to generate an in-phase and quadrature demodulated signal. The mixing signals have some phase difference ^^relative to the transmit signal, due to known imperfections in thesynchronisation of the transmission and mixing clocks. The mixer outputs for the in-phase^^^^^^ 709 and the quadrature ^^^^^^ 710, are expressed as:^^^^^^ = ^^^^^^ × ^^ cos^^^^ + ^^^ The nature of the receive signal of the receive transducer element 707 post LNA amplification 708, denoted by ^^^^^^can be simplified to the sum of three sinusoids, each representing a major signal source that contributes to the receive signal. The first sinusoid is the signal of interest, the Doppler signal. It is created when reflections occur at medium boundaries where the next medium (i.e. blood) moves relative to the transducer 707 The result of this moving boundary reflection is reflected sinusoid with a frequency that is shifted ^^from the original transmit frequency. This Doppler signal arrives back at a receiving transducer element with a random phase shift ^^due to the long path of the signal relative to the wavelength of the signal. The second type of signal comes from any ultrasound reflection of a medium that is stationary relative to the source. This results in no frequency shift, and so this signal has a return frequency of ^!with some random phase ^"due to its long path. The third signal originates due to direct mechanical coupling (crosstalk) between the transmitting and receiving transducer elements. This third signal has a frequency ^!with negligible phase shift relative to the original transmit signal, due to the close placement of the receiver and the transmitter transducer elements. The sum of these three signals makes the receive signal, each given an amplitude ^!(reflected transmit amplitude), ^^(reflected Doppler amplitude) and ^#(crosstalk amplitude). Since only a small fraction of signal power reflects off a moving mediumboundary, ^! ≫ ^^ and ^# ≫ ^^. Defining ^ = 0 as the moment of ultrasound wavetransmission, each sinusoid is multiplied by a unit step with some delay to represent the time it takes for the ultrasound signals to reflect and reach the receiver transducer element. Due to the close placement of the receiver transducer element relative to the transmitter transducer element (for example <1mm, optionally 0.6mm), the delay is negligible for the crosstalk signal. This receive signal is passed to the in-phase and quadrature mixers 709 and 710, generating ^^^^^^and Each sinusoid contained in the receive signal is multiplied by the corresponding in-phase or quadrature mixing sinusoid. Using the trigonometric identity. ^cos^^ + )^ + cos^^ − )^^The expression for the mixed signal is expanded, giving a sum of high frequency terms, Doppler frequency terms, and zero frequency (DC offset) terms. Below is an example for the in-phase signal, however similar steps can be applied to the quadrature. This signal is then low pass filtered (LPF) at 711, removing high frequency terms and admitting only the sinusoids containing the Doppler information and step content. Referring to Fig.14, the left plot (a) shows an arbitrary plot of the low pass filtered output, containing some Doppler and a singular DC step. The cancellation of the transmit frequencies during mixing gives rise to the DC offset terms in the above expression. The magnitude of the DC step 1401 in Fig 14. is random in nature, depending on several random factors such as the phase of the mixer signal relative to the transmitter ^^, the phase of the reflected transmit signal ^!, as well as their amplitudes. Due to the large amplitudes of the crosstalk and the reflected transmit signal, this voltage step magnitude is much larger than the amplitude of the doppler signal.1402 portrays the moment in time the signals arrive at the receiver, creating the DC voltage step component. The DC voltage step component is shown below. A mixer – transmitter phase locking circuit can be implemented to lower the DC step voltage magnitude by reducing the contribution from the crosstalk. This is done by settingthe mixer phase to be ^^ = , / 4, which minimises the crosstalk contribution to the DCvoltage step for both the in-phase and quadrature channel, since: To be able to sample the Doppler signal, the Doppler signal needs to be amplified large enough such that its amplitude can be resolved by an analogue-digital converter (ADC). As the Doppler signal is superimposed on top of the DC voltage step, amplifying the Doppler signal also amplifies the DC offset. Trace 1403 shows the small amplitude Doppler signal superimposed on the large magnitude DC step. The size of the DC step is much larger than the Doppler signal, so any attempt at amplifying the Doppler signal (whilst the DC step is present) saturates (or “clips”) any operational amplifier circuit with a gain large enough to enlarge the Doppler signal to a sampleable amplitude. Hence, the DC voltage step must be removed prior to any large amplification to avoid clipping. In traditional CW Doppler imaging, beamforming allows continuous transmission and receival of a transducer array, whilst the focus of the ultrasound beam is adjusted to scan across a region, in preferred embodiments to detect blood flow. As the channels are continuously active, the “step” nature of the signal can be ignored, and this DC voltage step can be instead considered as a constant DC offset in the signal. In this case, a simple high-pass filter is sufficient to remove the DC signal, with a pole in between zero frequency and the lowest typical Doppler frequency desired to be resolved. In the case of the portable ultrasound device 100 depicted in Fig. 1, the transmitter and receiver transducer element pairs are not continuously firing, but rather switching on, sampling for a short controlled dwell period (according to scan patterns as described above) albeit consistent with continuous wave doppler operation, and then switching off. In this case, the step in the DC voltage can no longer be ignored. The DC voltage step contains non- zero frequency content. The magnitude of this frequency content drops off with increasing frequency (Fourier transform of a step). Some of this frequency content will lie above the pole of a traditional high-pass filter and therefore be admitted. This manifests as a transient pulse in the high-pass filter output that occurs the moment after a pair of transducer elements are activated, distorting the Doppler signal temporarily. In traditional CW beamforming, the momentary nature of the pulse results in the pulse not being an issue given the long duration of beamform scanning. The right plot (b) of Fig.14 plots an arbitrary high pass filter response to the DC step natured signal shown on the left-hand side plot (a).1402 is the time instance the DC step signal occurs. The high pass filter admits the high frequency content of the DC step, and due to amplification by gain G, the step content is amplified to the operational amplifier saturation voltage 1404. The signal remains saturated for some time 1405, then begins to settle, eventually settling at the time instant 1406, which can be on the order of 100 milliseconds. Even if the pulse does not saturate, due to the overlapping admitted frequency content of the step and the Doppler signal, the Doppler signal cannot be differentiated from the pulse, and so the overall dwell time must be increased to include the settling time and the sampling time. An increase in transmitter transducer element dwell time proportionally increases scan time (recalling exemplary dwell times of 20 ms mentioned above), and so this effect dramatically hinders usability, as frames cannot be generated frequently enough to achieve sufficient scouting when a high-pass filter is used to remove the DC step voltage signal, referred to as “DC offset” in description below. The DOCU 105 provides a solution to this problem by allowing fast removal of the DC offset without the need of a high-pass filter which would reduce scanning rate and performance of the ultrasound device 100. The DOCU 105 first processes a CW mixer output signal with the CW mixer output processing circuit 701. In this embodiment, the CW mixer output signal of 701 is passed to the DOCU 105, where it is first prepared for DC offset removal using a differential to single ended amplifier 705 that has the function of converting the voltage differential output of the CW mixer to a single-ended signal. One embodiment of DC offset correction unit 105A uses a DC sample and hold circuit 702. In this embodiment, sampling of the output from 701 samples the DC offset in the analog domain by charging a hold capacitor. A switch is connected between differential amplifier 703 and the hold capacitor and is closed while charging the hold capacitor. After the hold capacitor is charged, the switch is opened and the hold capacitor drives the input of the next amplifier, a hold buffer amplifier. The hold buffer amplifier may be a unity gain buffer, and its purpose is to redrive the hold signal to the difference amplifier. More capacitors may be added to the circuit to reduce the effects of leakage current which results in a drift of the hold signal. To further reduce drift of the hold signal, the capacitors preferably have a dielectric absorption constant of less than 5%, more preferably less than 3%, even more preferably less than 1% and most preferably less than 0.1%. This circuitry has the advantage of removing the DC offset in a fast manner in comparison to other techniques, such as use of a high pass filter as previously practised in the art, resulting in an increased ultrasound scan rate. Another preferred embodiment of the DOCU is 105B as shown in Fig. 7B. Within this embodiment, the low pass filtered CW mixer outputs of 701 (derived from echo signal derived from transducer 707), as described above, are again converted to single ended signals by differential to single ended amplifiers 705. These amplifiers may buffer or provide a small amount of gain to the mixer output signal, amplifying the Doppler and DC offset by some amount but not to the extent that the output voltage saturates the operational amplifier output. This single ended signal is sampled by a single ended input analogue-digital converter (ADC) 712. The Doppler signal content is small relative to the DC offset, so the resultant digital signal approximates the magnitude of the DC offset content present in the mixer output signal. The DC sample circuit connects to processor 109. The processor 109 sends information to a DC recreation digital-analogue converter (DAC) 706, which recreates the sampled voltage magnitude, approximately recreating the DC offset. This DC offset is recreated once, and remains constant during the entire duration of a dwell period for a transmitter transducer element, say 20 ms dependent on the design scanning rate for the ultrasound device 100. Advantageously, the DC offset may be resampled and the DAC output voltage updated in the case where the DC offset value “drifts” causing clipping of the system. This may be advantageous in longer dwell period where the DC offset value can “drift” for longer. The digital-analogue converter 706 output may contain some noise content, and so its output can be filtered by a low pass filter (LPF) prior to passing to the next stage. The time constant of the low pass filter (LPF) is selected to minimise the admitted noise spectrum whilst ensuring the DC offset can reach the next stage within a reasonable time, through selection of a sufficiently short filter time constant. The recreated approximate DC offset magnitude is passed into one terminal of a differential operational amplifier circuit 703. The buffered (or slightly amplified) low pass filtered CW mixer output of 705 is input to the other input of the difference amplifier 703. The difference between the input terminals of the differential amplifier 703 equals the Doppler signal plus some small DC offset measurement error. The differential amplifier amplifies this signal, resulting in an output that contains the amplified Doppler signal plus the amplified DC offset measurement error. The gain of the differential amplifier is selected to ensure that the DC offset measurement error voltage will not cause saturation of the differential amplifier output voltage. The differential amplifier stage may be configured to be a low pass filter as well as an amplifier, to further attenuate high frequency noise on the output. The differential amplifier 703 output connects to a second analog to digital converter 704 which provides the isolated Doppler signal which is used in further processing to enable imaging of vasculature. An alternative DOCU architecture uses multiplexing to configure the analogue digital converter 704 to be the analogue to digital converter used in the DC sample ADC 712 to reduce the number of unique analogue-digital converter integrated circuits on the printed circuit board (PCB) and reduce space requirements for the DC offset removal circuit. During DC offset sampling, the multiplexers direct the single ended converted mixer output signal to the analogue-digital converter 704, where after digital-analogue DC offset recreation, the multiplexers connect the analogue-digital converter 704 to the difference amplifier 703 output for sampling of the Doppler signal. DOCU 105B removes DC offset in time close to 0 ms, very substantially less than a high pass filter (around 100 ms) which would compromise scanning rate. The processor 109 is configured to receive and process the isolated and amplified Doppler signal output from analogue-digital converter 704. During the dwell period of the transducer 707, the characteristics of the receive signals may fluctuate causing a ‘drift’ effect in the real DC offset value. This may result in a linear or non-linear ‘drift’ of the sampled isolated Doppler signal that may be mistaken as a Doppler shift in the frequency domain. In the present embodiment, the sample Doppler signal is prepared for analysis by utilising a linear fit to model the ‘drift’ which is then subtracted from the Doppler signal, to further isolate the Doppler signal. It is understood that other fitting models can be used dependent on the shape of the ‘drift’. The processor 109 is further configured to process components of a CWD signal with an example of such a signal being illustrated in Figs.8A and 8B. In the present embodiment, the components of the CWD signal contains the magnitude and phase difference, referred to simply as “phase”, of the quadrature (Q) and in-phase (I) channels for all frequencies present in the Doppler signal 804. Unless otherwise stated, in this specification, references to “phase” are to the phase difference between the I and Q channels. Fig.8A illustrates a magnitude-frequency plot 801 for a single transducer element where the received CWD magnitude signal is discretised to frequency bins 802. In one embodiment, the magnitude-frequency data points 804 give an indication of the presence of flow in a vessel, where a high magnitude across one or multiple frequency bins indicate flow for that transducer element. Fig.8B illustrates a phase-frequency plot 805 for a single transducer element where the received CWD phase difference signal is discretised to frequency bins 802. In one embodiment, the phase-frequency data points 804A provide an indication of the presence of blood flow, where one or multiple frequency bins having a phase of ±90° indicate blood flow for that transducer element. After conversion of the CWD signal to magnitude-frequency and phase-frequency data through a Fast-Fourier Transformation, processor 109 completes a magnitude FFT analysis 106 and phase FFT analysis 107. Fig.9A illustrates a flow chart that shows the steps involved in extracting values that represent the magnitude for the specific transducer element. As the magnitude-frequency CWD signal is from non-static tissue, movement interference could be present in the CWD signal. Additional sources of interference that could be present in the signal include electronic interference and acoustic interference. Each interference source has a distinctive magnitude-frequency pattern that could obscure the presence of blood flow and cause magnitude FFT analysis 106 to extract non-representative values. One example is slow tissue interference, such as tissue wall movement or small movements of the ultrasound device 100, where such interference presents low frequency spikes in the magnitude-frequency domain. Another example is fast tissue interference, such as large movements of CWD device 100, where such interference presents as high frequency spikes in the magnitude-frequency domain. The magnitude FFT analysis 106 commences with a magnitude smoothing algorithm 901 that reduces the previously mentioned interference present in the magnitude-frequency plot 801. In embodiments, the magnitude smoothing algorithm 901 may include smoothing filters, mean, median, Gaussian smoothing and Savitzky-Golay filters. These filter examples are not intended to be exhaustive. Following magnitude smoothing 901, the smoothed magnitude-frequency undergoes a magnitude extraction algorithm 902. In the present embodiment, the extraction algorithm 902 is a formula such as, but not limited to, the maximum magnitude value of the data, average magnitude value of the data or median magnitude value of the data. Conveniently, a prior step to the above listed formulae may include segmenting the data to focus on frequencies of interest using functions such as a step function or Butterworth filter. Advantageously, segmenting selected frequency ranges allows extracting a magnitude value from frequency ranges where flow characteristics might be present. One example involves detection of slow flow where a Butterworth or similar filter could be used to amplify the lower frequency magnitudes before extracting the maximum value. Advantageously, the lower frequency is where slow flow would be present. Alternatively, a step function could be applied to amplify higher frequencies, where fast flow would be present. The extracted magnitude values 906 are then a set of input data to the reconstruction algorithm 108. Fig. 9B illustrates a flow chart that shows the steps involved in extracting values that represent the phase for the activated transducer element. As the phase-frequency CWD signal 905 is from non-static tissue, movement interference could be present in the signal. For example, vascular structures have a motion which influences the signal and such interference is compensated for here. Additional sources of interference that could be present in the signal include electronic and acoustic interference. Similar to the magnitude-frequency domain, each interference source has a distinctive phase- frequency pattern that could obscure the presence of blood flow and cause phase FFT analysis 107 extracting non-representative values. Examples are slow tissue interference, such as tissue wall movement or small movements of the ultrasound device 100, where both noise sources present as phase values not in the range of ±90° in the phase-frequency plot 805. The phase FFT analysis 107 begins with a phase smoothing algorithm 903 that reduces the previously mentioned interference present in the phase-frequency plot 805. In the present embodiment, a suitable phase smoothing algorithm includes – but is not limited to – smoothing filters, mean, median and Gaussian smoothing. Following phase smoothing 903, the smoothed phase-frequency is processed by a phase extraction algorithm 904. In the present embodiment, this extraction algorithm 904 is a formula such as – but not limited to – maximum phase value of the data, average phase value of the data or phase magnitude value of the data. Conveniently, a prior step to the above listed formulae may include segmenting the phase data to focus on selected frequencies of interest using functions such as a step function. Advantageously, segmenting selected frequency ranges of interest allows extracting phase values from frequency ranges where flow characteristics might be present. An example is the use of a step function to threshold low or high frequencies depending on if the desired blood flow speed detected is slow or fast. Alternatively, prior to extraction, the data may also be segmented for the frequency or frequencies that were extracted from the magnitude extraction algorithm 902. The extracted phase values 907 are then a set of input data to the reconstruction algorithm 108. The processor 109 is further configured to process the extracted magnitude and / or phase values (preferably both) using a reconstruction algorithm 108 to display a representation of identified vasculature flow, for example in a candidate vein for a cannulation procedure. An example of such a reconstruction algorithm is shown in outline in Fig.10A. In this embodiment of reconstruction and visualisation, processor 109 processes an array of magnitude and phase values where each value of the array is extracted using the extraction algorithms described with reference to Figs. 9A and 9B for each transducer element. The dotted line around the phase / magnitude mapping stage 1001 represents that this stage of processing is optional and is completed if both magnitude and phase data are present. The aim of phase / magnitude mapping stage 1001 is to tie together the flow information present in both magnitude and phase data with the objective of increasing values for transducer elements that are indicated to be above vasculature of interest, for example a potential vein for cannulation. One embodiment of phase / magnitude mapping is the product of the magnitude, which may be normalised, with a function of the corresponding phase. The following are a list of example functions that may be used with phase, where phase is θ: 1906 − 78 < : ≤ 906 +2^3^ = 4 0 '^ℎ=>?@(=−1 −906 − 78 < : ≤ −906 where a1 is preferably less than 60°, more preferably less than 45° and most preferably less than 30°: where b1 is preferably 90° f(x) = {sinn(θ) -180° ≤ θ ≤ 180°} where n is preferably positive odd integer values. Additionally, the resulting output from phase / magnitude mapping may undergo a further transformation through a tan function, in order to amplify flow signals. Alternatively, only the magnitude or phase values are used for the subsequent processing steps. The function of the processing step array plot processing 1002 is to minimise additional interference, not reduced during extraction as described with reference to Figs.9A and 9B. In the present embodiment, this includes – but is not limited to – smoothing filters, mean, median and Gaussian smoothing. Advantageously, this stage may include a simple or adaptive thresholding to further eliminate transducer elements that have interference. An example of such thresholding is setting any values in the data array to 0 for values that are not indicative of blood flow where the thresholding value is determined by, but not limited to, a percentage of the maximum value in the array, a predetermined set value or the mean of the values. The purpose of the array display function 1003 is to display and visualise the processed data set as an image on the display 110 of portable ultrasound device 100 in a straightforward manner with a higher quality signal processing method resulting in a stronger, more stable, lower interference signal at faster speed than previously described in the case of the Applicant’s International Publication WO 2022221913 incorporated herein by reference. A reconstruction algorithm enables the extracted data to be presented as an image or visualised on display 110. Preferred embodiments are described below though it will be understood that other reconstruction algorithms could be used. The terms “reconstruction algorithm” and “flow reconstruction algorithm” are used interchangeably. Fig. 10B provides a display of the data acquired from the reconstruction process as described above with reference to Fig.10A. Each line 1004 is a line graph of the values for each transducer array of ultrasound device 100, in this case having three transducers with the peaks observed being representative of blood flow detected under the transducer elements of each transducer array. In another embodiment of reconstruction algorithm 108, outlined in Fig. 11A, the processor 109 is configured to process the extracted magnitude values 906 and / or phase values 907 (preferably both 906 and 907) to display a line representative of the underlying vasculature through which blood flow has been detected during transducer scanning. Similarly to the embodiments of Figs.10A and 10B, processor 109 processes an array of magnitude and phase values where each value of the array is extracted using the process described above with reference to Figs.9A and 9B for each activated transducer element. The function of the peak detection algorithm 1101 is to determine, from the magnitude data, which subset of transducer elements of the array has underlying blood flow. This can be completed using thresholding, where transducer elements that have a magnitude over a predetermined or adaptive threshold are considered transducer elements with underlying blood flow indicative of presence of a vessel through which that blood flows. Alternatively, common peak techniques, such as derivatives, wavelets and / or convolutions may be used to detect peaks across a transducer to determine transducer elements with underlying flow. It is understood that the starting data is not exclusive to magnitude and includes any data that may give an indication that flow is present in the specific transducer element, for example phase and time domain. A binary dataset of transducer elements with flow and no flow is outputted to the next process step, line connection algorithm 1102. Line connection algorithm 1102 processes both the flow dataset and extracted phase values 907 for each transducer element to connect lines across the transducers indicating where a vasculature is present. This is accomplished by matching peaks that are likely to be from the same vessel, where matching is determined based on characteristics such as phase, magnitude, location and other acoustic or physical properties. Advantageously, transducer elements located over the same blood vessel would have similar values for the selected characteristic. For example, a group of transducer elements from a peak of one transducer may have similar phase values to a group of transducer elements from a peak of an adjacent transducer, resulting in a match between the two peaks on the basis of the similar phase values as determined by processor 109. In another example, a group of transducer elements from a peak of one transducer may have a similar position to a group of transducer elements from a peak of an adjacent transducer, resulting in a match between the two peaks. In yet another example, a group of transducer elements from a peak of one transducer that has been identified as a vein may be matched to a group of elements from a peak of an adjacent transducer that is also identified as a vein. In another example, a group of transducer elements from a peak of one transducer may have a similar number of transducer elements in the group to a group of transducer elements from a peak of an adjacent transducer, resulting in a match between the two peaks. In another example, a group of transducer elements from a peak may have similar vessel characteristics to a group of transducer elements from a peak of an adjacent transducer, including but not limited to vessel diameter, vessel velocity, vessel depth and vessel ellipticity. The vessel characteristics may be measured using other forms of ultrasound such as A-mode or B-mode. In another example, a group of elements from a peak of one transducer may have any combination of the above described similarities to a group of transducer elements from a peak of an adjacent transducer, resulting in a match between the two peaks. Matched transducer elements would be considered to be from the same blood vessel and a line is drawn between the transducer elements of the parallel arrays of transducer elements. The purpose of line display function 1103 is to display and visualise the processed dataset to the display 110 of portable ultrasound device 100. Fig.11B is an example of the visual display from the data acquired from the reconstruction algorithm as outlined in Fig.11A. In this embodiment, ultrasound device 100 has three transducers and detected blood vessels are represented as lines 1106 between transducer elements 1104, 1105 that were matched by the line connection algorithm 1102. The difference in shade between the left dots 1104 and right dots 1105 represents that each line shows different vascular types (left is venous and right is arterial). This discrimination of vascular type (i.e. vein or artery) can be inferred from flow direction, which is done by line connection algorithm 1102, where transducer elements can be matched for similar phase values in ranges that represent away (-90°±X°) or towards flow (90°±X°) where X is preferably less than 60°, more preferably less than 45° and most preferably less than 30°. The different vasculature may be represented by different colours, for example blue for veins and red for arteries or any desired colour or colour combination. The line display function 1103 has the advantage, in comparison to an array display, of displaying a line which requires less training from the user to determine where the underlying vessel is located. However, the processing requirements are higher in comparison. In yet another embodiment of the reconstruction algorithm 108, outlined in Fig.12A, the processor 109 is configured to process the extracted magnitude 906 and or phase 907 values (preferably both 906 and 907) to display a heatmap representative of the underlying vasculature. Similarly to the previous embodiment, processor 109 processes an array of magnitude and phase values where each value of the array is extracted using the process described in Figs. 9A and 9B for each transducer element. The phase / magnitude mapping algorithm 1001 is the same as outlined in Fig.10A. Similarly, the phase / magnitude mapping algorithm is optional and the reconstruction may proceed with just magnitude or phase values. The objective of the heatmap colour algorithm 1201 is to apply a colour to each transducer element dependent on the strength of detected blood flow underneath that transducer element. The algorithm used is dependent on the input data that is used. In one embodiment, where magnitude is the only input data, the following algorithm may be applied. The magnitude array data is normalised by the maximum value and colours are assigned to represent no flow (minimum value of the normalised data) and flow (maximum value of the normalised data). Values in the normalised data that are between the minimum and the maximum are assigned a colour based on a linear interpolation between the no flow colour and flow colour in the RGB or HSV space. The colour of no flow is preferably black; however, any colour could be used. The colour of flow is preferably blue, however any colour may be used including colours that are typically associated with vasculature (e.g. purple or red). The result is transducer elements that have the maximum flow may be blue, transducer elements with weaker flow may be darker blue and transducer elements with no flow may be black. In another embodiment of the heatmap colour algorithm 1201 where phase is the only input data, the following algorithm is applicable. Colours are assigned to the phase data to represent towards flow (90°), away flow (-90°) and no flow (0°, ±180°). Preferably, the colours are blue, red and black respectively following a convention as described above. However, any combination of colours may be used. Values in the phase data that are between these values are assigned a colour based on the following function: QMNTSLMNTSQMNO#P 906 < : ≤ 1806where X is a linear interpolation between the two stated colours in the RGB or HSV space. The result is elements that have the maximum ‘towards’ flow will be blue, elements with the maximum ‘away’ flow will be red, elements with weaker ‘towards’ or ‘away’ flow will be a darker blue or red and elements with no flow will be black. In yet another embodiment of the heatmap colour algorithm 1201, where the processed phase / magnitude mapped dataset from the phase / magnitude mapping algorithm process 1001 is the input data, the following algorithm may be applied. The phase / magnitude mapped dataset is normalised with the maximum absolute value such that the strongest towards flow will be a value of 1, away flow will be a value of -1 and no flow will be a value of 0. Similar to the previous described algorithms, each value is assigned colour, preferably blue, red and black respectively. Values in the normalised phase / magnitude mapped dataset that are between these values are assigned a colour based on the following function: where X is a linear interpolation between the two stated colours in the RGB or HSV space and c1 is preferably less than 0.20, more preferably less than 0.15 and most preferably less than 0.10. The result is transducer elements that have the maximum towards flow will be blue, transducer elements with the maximum away flow will be red. Transducer elements with weaker towards or away flow will be a darker blue or red and transducer elements with no flow will be black. It is to be understood the x range may be further segmented and more colours included to the output. The purpose of the heatmap display function 1202 is to display and visualise the processed dataset to the display 110 of portable ultrasound device 100. Fig.12B is an example of the heatmap visual display from the data acquired from the reconstruction algorithm as outlined with reference to Fig.12A. For this embodiment, three transducer arrays are again present, and the presence of underlying vasculature may be inferred from the coloured heat map, where the vessel is on an angle from the middle of the bottom transducer to the left of the top transducer. Transducer elements are displayed as their assigned colour from the heatmap colour algorithm 1201. In this embodiment, the darker shade element 1203 would be indicative of strong blood flow and would either be a red or blue dependent on flow direction. The lighter shade elements 1204 are representative of a weaker blood flow in comparison and would be the same colour as the darker shade elements 1203 but at a different shade. Additional to the visualisation of the processed dataset, the heatmap display function 1202 includes an interpolation algorithm to increase the number of displayed transducer elements allowing for a smoother displayed image. This interpolation algorithm may, without limitation, be selected from the group consisting of bilinear, bicubic, inverse distance or multi-dimensional spline algorithms. In yet another aspect, the processor 109 is further configured to perform reconstruction based on the time domain Doppler signal rather than performing reconstruction based on the frequency domain Doppler signal. In one embodiment relating to processing and analysis in the time domain, one of the global or local voltage peak-to-peak (Vpp) value of the Doppler signal for both the I and Q channels can be used as an indication of blood flow. Doppler signals with low Vpp are correlated with no blood flow detected and signals with high Vpp are correlated with blood flow detected. Advantageously, this methodology is still sensitive to blood flow in cases where a full period of the desired Doppler shift is not captured with the trade-off of losing vessel flow direction information. This time domain analysis technique can replace or be utilised in conjunction to both the magnitude FFT analysis 106 and phase FFT analysis 107 and become an input to the vessel reconstruction algorithm 108. For example, in the time domain reconstruction algorithm described in Fig 10A, extracted Vpp values may undergo Vpp mapping through normalisation with the global max Vpp or transducer max Vpp. Alternatively, the following normalisation formula can be used to reduce values close to the noise floor: where 7 is a constant that is preferably close to the average Vpp of the slowest desired flow for candidate vasculature, in particular a vein for cannulation. This stage may include a simple or adaptive thresholding to further eliminate transducer elements that have interference. An example of such thresholding is setting any values in the data array to 0 for values that are not indicative of blood flow where the thresholding value is determined by, but not limited to, a percentage of the maximum value in the array, a predetermined set value or the mean of the values. These values would then be processed by array display function 1003. Likewise the time domain may also be used in the methodology presented in Figure 11A where a peak detection algorithm is used to determine peaks of high Vpp that is indicative of flow. Then the line correction algorithm 1102 processes the Vpp data set and connects peaks that are likely from the same vessel. Alternatively, the time domain analysis may be completed in addition to the frequency domain analysis and become an additional extracted value set into the line connection algorithm 1102. The Vpp values are then an additional value set to match peaks, where peaks with similar Vpp values are most likely representative of the same vessel. In a similar manner, the time domain may also be used in the methodology presented in figure 12A as an input to the heatmap colour algorithm 1201 instead of or in conjunction with the frequency domain values. The normalised Vpp data may undergo colour mapping to apply a colour to each transducer element depending on the strength of detected blood flow underneath that transducer element. Since no flow direction information is discerned, a single colour gradient is used. where X is linear interpolation between the two stated colours. Black and white are used to convey that no direction information is provided, however it is understood that other combinations of colours may be utilised. Throughout this specification, unless the context requires otherwise, the word "comprise" or variations such as "comprises" or "comprising", will be understood to imply the inclusion of a stated integer or group of integers but not the exclusion of any other integer or group of integers

Claims

CLAIMS 1. A portable ultrasound device for non-invasively imaging a selected sub-cutaneous structure in a subject, comprising: (a) a housing; (b) a plurality of arrays of transducer elements, each array being obliquely angled and arranged in parallel and each transducer element comprising a transmitter transducer and a receiver transducer, located within said housing for continuously transmitting acoustic waves at ultrasonic frequency in a carrier signal having a predetermined frequency range toward a body of a subject and continuously receiving echo signals in a predetermined frequency range from the body of the subject following reflection of ultrasound energy, said plurality of arrays of transducer elements enabling imaging of the sub-cutaneous structure in multiple transverse and lateral planes; (c) an acoustic lens disposed relative to said plurality of arrays of transducers for manipulating transmission of said acoustic waves; (d) a transducer controller for operating said plurality of arrays of transducer elements in continuous wave doppler mode and communicable with a processor, operable under control of electronic program instructions, for processing said received echo signals from said plurality of arrays of transducer elements, wherein said transducer controller is operable, under control of the electronic program instructions to control a scanning order and scanning rate of a plurality of transducer elements within said plurality of arrays of transducer elements for generating a Doppler signal processable by the processor to provide an image of the sub-cutaneous structure.

2. The ultrasound device of claim 1, wherein said parallel arrays are angled at an angle of insonation Φ where 45<Φ<80°.

3. The ultrasound device of claim 1 or 2, further comprising at least one continuous wave mixer for mixing the carrier signal with the received echo signals to form amixer output signal, the processor isolating a Doppler signal from the mixer output signal, wherein the processor executes a sequential scan pattern.

4. The ultrasound device of claim 3, wherein said sequential scan pattern comprises activating a first transducer element for a dwell period sufficient to capture a Doppler signal, deactivating the first transducer element at the end of the dwell period and activating a second transducer element for the dwell period.

5. The ultrasound device of claim 4, wherein said sequential scan pattern is repeated until all of the transducer elements have been activated for the dwell period.

6. The ultrasound device of claim 3, wherein a plurality of transducer elements, optionally a pair of transducer elements, are activated in parallel with the receive signals being input to the continuous wave mixer with Doppler signals from the plurality of transducer elements.

7. The ultrasound device of claim 6, wherein more receive transducer elements are activated than transmit transducer elements during a scan.

8. The ultrasound device of any one of claims 4 to 6, wherein the transducer controller controls the scanning rate, optionally by controlling the dwell period.

9. The ultrasound device of claim 3, comprising a plurality of mixers and the processor concurrently activates each of a selected plurality of transducer elements in a scan pattern executed by the processor.

10. The ultrasound device of claim 9, wherein the selected plurality of transducer elements has the same number of transducer elements as the number of mixers, the processor activating the selected plurality of transducer elements for a dwell period sufficient to capture a Doppler signal.

11. The ultrasound device of claim 8, wherein each received echo signal is processed by a dedicated mixer.

12. The ultrasound device of any one of claims 3 to 11, wherein the processor isolates the Doppler signal from the mixer output signal by subtracting a DC offset signal variable with each transducer element activation from the mixer output signal.

13. The ultrasound device of any one of claims 7 to 12, wherein the selected plurality of transducer elements excludes physically adjacent transducer elements.

14. The ultrasound device of claim 13, wherein the activated transducer elements are separated by at least one inactive transducer element.

15. The ultrasound device of claim 3, comprising a single mixer wherein the processor activates a plurality of physically adjacent transducer elements.

16. The ultrasound device of claim 3, wherein the processor activates each of a first group of transducer elements in a first sequential scan pattern and, on detection of a selected sub-cutaneous structure by processing of the received echo signals from a second group of transducer elements comprised within said first group of transducer elements, activates transducer elements within said second group of transducer elements in a second sequential scan pattern, optionally wherein the second group of transducer elements has a lesser number of transducer elements than the first group of transducer elements.

17. The ultrasound device of claim 16, configured to enable switching between the first group of transducer elements and the second group of transducer elements.

18. The ultrasound device of claim 16 or 17, including a switch for each of the groups of transducer elements to allow control over field of view by changing the number of activated transducer elements.

19. The ultrasound device of any one of the preceding claims, wherein the transducer controller enables an adaptive scan pattern utilising a plurality of scan protocols, each scan protocol containing an activation configuration and a corresponding flow reconstruction.

20. The ultrasound device of claim 19, wherein the transducer controller uses a first scan protocol and a second scan protocol, the first scan protocol having anactivation configuration with a short dwell period, optionally 20 ms, optimised for scan rate and a second scan protocol having an activation configuration with a relatively longer dwell period optimised for displaying a blood vessel, optionally with flow direction information.

21. The ultrasound device of claim 20, configured to enable a user to switch between the plurality of scan protocols, optionally through a user interface.

22. The ultrasound device of claim 20 or 21, configured to control switching between the plurality of scan protocols wherein transducer elements detecting flow from the first protocol are used as a transducer element mask for the second protocol to increase scan rate of the second protocol while preserving vessel information.

23. The ultrasound device of any one of the preceding claims, comprising a user interface enabling a user to command the transducer controller to change field of view, optionally through enabling the transducer controller to activate a first group of transducer elements in a first sequential scan pattern and activate a second group of transducer elements in a second sequential scan pattern.

24. The ultrasound device of claim 23, wherein the first sequential scan pattern is in one dimension and the second sequential scan pattern is in two dimensions.

25. The ultrasound device of claim 24, wherein the user interface allows the user to switch between a one dimensional view and a two dimensional view.

26. The ultrasound device of any one of claims 23 to 25, wherein the processor compares the received echo signal from the first sequential scan pattern with the received echo signal from the second sequential scan pattern to determine if the ultrasound device has moved.

27. The ultrasound device of claim 26, wherein, on determination that the ultrasound device has moved, the processor activates a third group of transducer elements, the third group of transducer elements containing a greater number of transducer elements than the second group of transducer elements.

28. The ultrasound device of claim 26, wherein, on determination that the ultrasound device has not moved, the processor activates a third group of transducer elements, the third group of transducer elements containing a lesser number of transducer elements than the second group of transducer elements.

29. The ultrasound device of any one of claims 2 to 28, wherein received echo signals from the receiver transducer elements are mixed with the carrier signal by said at least one mixer to remove the carrier signal by quadrature modulation and generate an output signal, where the result for the I-channel and Q-channel after removal of the high frequency components and comprising a DC offset value is the following equation:

30. The ultrasound device of claim 29, wherein the processor removes the DC offset to isolate the Doppler signal without a high pass filter.

31. The ultrasound device of claim 30, wherein the processor includes circuitry to remove the low frequency DC offset from the output signal, the circuitry comprising: a first sample and hold circuit for measuring the DC offset signal, hold the DC offset signal and redrive the hold signal; and a difference amplifier that subtracts the measured DC offset signal from the mixer output signal to isolate the Doppler signal.

32. The ultrasound device of claim 12 or 30, wherein the processor samples the output signal for each activation of transducer element(s) to separately determine the DC offset signal for each said activation, determines the DC offset signal for eachsample and subtracts the DC offset signal from the output signal to isolate the Doppler signal.

33. The ultrasound device of claim 32, wherein the processor is configured to determine drift in DC offset signal as a function of fluctuation in received echo signals, the drift being subtracted from the Doppler signal to further isolate the Doppler signal.

34. The ultrasound device of any one of the preceding claims, wherein the processor processes at least one of the following components of the generated Doppler signal to image the sub-cutaneous structure: a digital raw voltage Doppler signal, a magnitude-frequency spectrum of the Doppler signal, the phase-frequency spectrum of the Doppler signal.

35. The ultrasound device of claim 34, wherein the processor processes the magnitude- frequency spectrum of the generated Doppler signal.

36. The ultrasound device of claim 35, wherein the processor processes both the magnitude-frequency spectrum and the phase-frequency spectrum of the generated Doppler signal.

37. The ultrasound device of claim 35 or 36, wherein the processor processes the magnitude-frequency component to extract magnitude values that are representative of the sub-cutaneous structure.

38. The ultrasound device of claim 37, wherein the magnitude-frequency component is processed by a magnitude-Fast Fourier Transformation (FFT) method.

39. The ultrasound device of any one of claims 34 to 38, wherein the processor processes the phase-frequency component to extract phase values that are representative of the sub-cutaneous structure.

40. The ultrasound device of claim 39, wherein the phase-frequency component is processed by a magnitude-Fast Fourier Transformation (FFT) method.

41. The ultrasound device of any one of claims 37 to 40, wherein, prior to processing of component(s) of the Doppler signal, interference is removed from the Doppler signal.

42. The ultrasound device of any one of claims 37 to 41, wherein the processor executes an extraction algorithm to filter the Doppler signal to extract signals in the frequency or time domain.

43. The ultrasound device of claim 42, wherein the processor executes a reconstruction algorithm to process said signals at selected frequency to image the sub-cutaneous structure.

44. The ultrasound device of claim 42 or 43, wherein the processor executes a reconstruction algorithm to process said signals in the time domain to image the sub-cutaneous structure, optionally by using extracted peak to peak voltage (Vpp) values in the Doppler signal where the processor processes the Vpp data set to connect peaks that are likely from the same vessel.

45. The ultrasound device of claim 43 or 44, wherein the processor is configured to display the sub-cutaneous structure as an array corresponding with transducer elements detecting a signal from the sub-cutaneous structure.

46. The ultrasound device of claim 45, wherein the reconstruction algorithm maps extracted values of both magnitude and phase with a resulting mapped dataset or individual components of the signal being processed to display the sub-cutaneous structure as the array.

47. The ultrasound device of any one of claims 43 to 46, wherein the processor is configured to display the sub-cutaneous structure as a line linking transducer elements detecting a signal from the sub-cutaneous structure.

48. The ultrasound device of any one of claims 43 to 47, wherein the reconstruction algorithm detects peaks corresponding with the sub-cutaneous structure from the extracted values from the magnitude component, at least one group of peaks being formed based on probability that each peak is from the sub-cutaneous structure based on characteristics extracted from at least one of the time domain, frequencydomain and vessel characteristics the processor processing the at least one group of peaks to generate a line linking the groups of peaks and displaying the sub- cutaneous structure.

49. The ultrasound device of any one of claims 43 to 46, wherein the processor displays the sub-cutaneous structure as a colour map.

50. The ultrasound device of any one of claims 43 to 49, wherein the processor is configured to allow a user of the device to switch between reconstruction algorithms and visualisation modes.

51. The ultrasound device of any one of the preceding claims, wherein said sub- cutaneous structure is a vascular structure, optionally a vein.

52. The ultrasound device of claim 51, wherein the processor displays a coronal view of the vascular structure.

53. The ultrasound device of claim 51, wherein the processor displays one dimensional data representative of the vascular structure.

54. The ultrasound device of any one of claims 51 to 53, wherein the processor determines the direction of flow in the vascular structure.

55. The ultrasound device of any one of the preceding claims, wherein the acoustic lens covers said plurality of arrays of transducer elements, preferably terminating in a flat surface at a lens tissue interface.

56. The ultrasound device of any one of the preceding claims, wherein the acoustic lens comprises a plurality of layers of materials, each layer being different and having at least one of different shape and different acoustic properties.

57. The ultrasound device of claim 56, wherein the acoustic lens comprises two layers, a first layer being a material in a convex shape and the second layer being a material having different acoustic properties to said first layer, said second layer providing an interface between the first layer and underlying tissue.

58. The ultrasound device of any one of the preceding claims, wherein the acoustic lens has substantially the same acoustic properties as human tissue 59. A method for imaging a sub-cutaneous structure in a subject with an ultrasound portable ultrasound device comprising: (a) a housing; (b) a plurality of arrays of transducer elements, each array being obliquely angled and arranged in parallel and each transducer element comprising a transmitter transducer and a receiver transducer, located within said housing for continuously transmitting acoustic waves at ultrasonic frequency in a carrier signal having a predetermined frequency range toward a body of a subject and continuously receiving echo signals in a predetermined frequency range from the body of the subject following reflection of ultrasound energy, said plurality of arrays of transducer elements enabling imaging of the sub-cutaneous structure in multiple transverse and lateral planes; (c) an acoustic lens disposed relative to said plurality of arrays of transducers for manipulating transmission of said acoustic waves; (d) a transducer controller for operating said plurality of arrays of transducer elements in continuous wave doppler mode and communicable with a processor, operable under control of electronic program instructions, for processing said received echo signals from said plurality of arrays of transducer elements, wherein said transducer controller is operable, under control of the electronic program instructions to control a scanning order and scanning rate of a plurality of transducer elements within said plurality of arrays of transducer elements for generating a Doppler signal processable to provide an image of the sub-cutaneous structure.

60. A method for imaging a sub-cutaneous structure in a subject with a portable ultrasound device, said method comprising: (a) continuously transmitting acoustic waves at ultrasonic frequency in a carrier signal having a predetermined frequency range toward a body of a subject from a plurality of arrays of transducer elements, each array being obliquely angled and arranged in parallel and each transducer element comprising a transmitter transducer and a receiver transducer;(b) manipulating transmission of said acoustic waves; (c) continuously receiving echo signals from the body of the subject following reflection of ultrasound energy; (d) operating said plurality of arrays of transducer elements in continuous wave doppler mode; and (e) processing said received echo signals from said plurality of arrays of transducer elements, wherein a scanning order and scanning rate of a plurality of transducer elements within said plurality of arrays of transducer elements is controlled for generating a Doppler signal processable to remove interference and provide an image of the sub-cutaneous structure.

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