Real-time photoacoustic and ultrasound imaging system and method

Methods and system for producing combined photoacoustic/ultrasonic image frames use a low-power narrow beam laser to direct sequential pulses along a path overlying an internal region of interest. Photoacoustic responses are received and used to generate sub-frames. Between each of the laser pulses a plurality of ultrasound pulse-echo beams are sequentially emitted towards the region of interest, and the reflections are received and used to generate ultrasound sub-frames. The photoacoustic sub-frames are combined to produce a photoacoustic frame, and the ultrasound sub-frames are combined to produce an ultrasound frame. The photoacoustic and ultrasound frames are combined to produce an image frame. The method and system are suitable for producing real-time, high-contrast video.

BACKGROUND

Photoacoustic imaging complements clinical ultrasound by adding optical absorption as a contrast mechanism at spatial resolutions comparable to ultrasound for penetration depths up to a few centimeters. Using the same transducers and imaging electronics, it can provide molecular information absent from ultrasound images. Ultrasound-guided photoacoustic imaging can have several clinical applications related to imaging the vasculature (exploiting hemoglobin contrast); detecting molecularly targeted, nanoscale contrast agents; and guiding interventional procedures. An integrated scan format in which photoacoustic/ultrasound image frames are interleaved at real-time rates is needed for optimal translation of this technology to the clinic.

Integrated photoacoustic/ultrasound systems have been developed as a research tool. Typically, high-pulse-energy (tens to hundreds of millijoules per pulse) Q-switched lasers are required to produce high signal-to-noise ratio (SNR) photoacoustic signals induced by nanosecond pulses. However, these laser sources are generally not cost effective, massive (usually >1 m long with a bulky power unit), and difficult to maintain. In addition, their low pulse repetition frequency (PRF) can severely limit image frame rates for real-time operation, hindering integration of photoacoustics with a real-time commercial ultrasound scanner for clinical applications. To facilitate clinical translation of interleaved photoacoustic/ultrasound imaging, a low-cost and portable laser is integrated with a commercial ultrasound scanner.

For example, portable fiber-amplified semiconductor-based lasers and laser diode-pumped Q-switched lasers are relatively inexpensive and output nanosecond pulses with repetition rates up to several hundreds of kilohertz. However, these lasers are not suitable for conventional photoacoustic imaging because the single pulse energies produced are quite low, in most cases from hundreds of microjoules up to a few millijoules. Such low energies distributed over an entire image volume of interest cannot generate photoacoustic signals with acceptable SNR characteristics for single-pulse operation. To produce acceptable optical fluence (energy per unit area) for high-SNR measurements, the beam must be focused to a small spot. This means the lateral extent of the image is quite limited. In addition, optical scattering is very strong for most biological tissues in the visible and near-infrared optical wavelength range, limiting light penetration, especially for a small diameter beam spot.

The SNR can be improved by averaging. However, the SNR improvement from averaging only increases as the square root of the number of pulses, and the PRF is limited by the acoustic time-of-flight to avoid range ambiguities (i.e., sources from multiple ranges delivering photoacoustic signals to the ultrasound array at the same time). For example, to image a 1-cm region, the laser pulse repetition interval (PRI) should be larger than the acoustic propagation time,

The maximum PRF is therefore approximately 154 kHz. Furthermore, laser safety regulations (maximum average power of 1 W/m2) also limit the repetition rate for a given laser fluence. For example, the maximum fluence allowed for a 1-kHz laser is 1 mJ/cm2, a value producing low optical intensities deep inside the body. Therefore, there is a tradeoff between SNR (related to laser fluence) and image frame rate (related to PRF).

As disclosed herein, these limitations may be overcome in part by combining the photoacoustic response of multiple laser pulses that are scanned over a region of interest. A high-repetition-rate (e.g., ˜1 kHz), low-energy (e.g., a few millijoules) laser enables rapid scanning Photoacoustic signals induced by each laser shot are recorded to form a sub-frame of radio frequency (RF) data. Photoacoustic sub-frames corresponding to the narrow beam sub-frame can be reconstructed to show regional information near the laser scanning spot. The final photoacoustic image covering the complete scan area is synthetically reconstructed by combining all RF sub-frames. A real-time interleaved photoacoustic/ultrasound system with a frame rate of 30 Hz or greater is disclosed.

One or more of the present inventors have demonstrated photoacoustic imaging using beam scanning, as described in more detail in J. Xia, C.-W. Wei, I. M. Pelivanov, and M. O'Donnell, “Photoacoustic imaging using narrow beam scanning,” in IEEE Int. Ultrasonics Symp., 2011, pp. 2380-2383, and in J. Xia, C.-W. Wei, L. Huang, I. M. Pelivanov, and M. O'Donnell, “Comparison of PA imaging by narrow beam scanning and one-shot broad beam excitation,” Proc. SPIE, vol. 7899, Art. No. 78991L, 2011, each of which is hereby incorporated by reference in its entirety.

SUMMARY

A method for generating a combined photoacoustic and ultrasound image frame of an internal region of a body includes (i) irradiating a first location on the surface of the body with a laser pulse to illuminate the internal region, and using the corresponding photoacoustic response to generate a first photoacoustic sub-frame; (ii) sequentially transmitting ultrasound pulse-echo beams towards a first section of the internal region, and using the ultrasound reflections to generate a first set of ultrasound sub-frames of the internal region; (iii) irradiating a second location on the surface of the body with a laser pulse to illuminate the internal region, and using the corresponding photoacoustic response to generate a second photoacoustic sub-frame; (iv) sequentially transmitting ultrasound pulse-echo beams towards a second section of the internal region, and using the ultrasound reflections to generate a second set of ultrasound sub-frames of the internal region; combining the first and second photoacoustic sub-frames to generate a photoacoustic frame of the internal region, and combining the first and second sets of ultrasound sub-frames to generate an ultrasound frame of the internal region; and (v) combining the photoacoustic frame and the ultrasound frame to generate an image frame of the internal region of the body.

In an embodiment the laser pulses have energies of less than 10 mJ, and in embodiment the laser pulses have energies of less than 2 mJ. In an embodiment the laser pulses have a diameter of less than 3 mm.

In an embodiment the laser pulses have a wavelength between 400 and 1500 nm, and a fluence less than a wavelength-dependent fluence that would damage tissue. In an embodiment the fluence is less than 100 mJ/cm2.

In an embodiment at least six ultrasound pulse-echo beams are directed towards the internal region after each laser pulse.

In an embodiment the laser pulses are produced by a laser diode-pumped Q-switched laser, and the ultrasound pulse-echo beams are generated with an ultrasound linear array that also receives the ultrasound reflections and the photoacoustic responses.

In an embodiment the ultrasound linear array is disposed in a probe, and the laser pulses are emitted from a distal face of the probe. The laser pulses may be transmitted to the probe by optical fibers extending along either side of the ultrasound linear array. In an embodiment the laser pulses are distributed from the laser to the optical fibers with a remote switching unit.

In an embodiment the methods described above are used to generate a real-time video of the internal region.

A method for generating a combined photoacoustic and ultrasound image of an internal region of a body includes irradiating the external surface of the body with sequential narrow-beam, lower power laser pulses that irradiate the internal region of the body, wherein the sequential laser pulses are directed to a plurality of different locations along a path on the external surface, and receiving photoacoustic signals from the body corresponding to the sequential laser pulses; between each of the sequential laser pulses, sequentially transmitting a plurality of ultrasound pulse-echo beams to the internal region of the body, and receiving ultrasound reflections of the plurality of ultrasound pulse-echo beams; generating a first set of sub-frames from the received photoacoustic signals; generating a second set of sub-frames from the received ultrasound reflections; and combining the first set of sub-frames with the second set of sub-frames to generate a combined photoacoustic and ultrasound image frame of the internal region of the body.

An imaging system for providing real-time images of an internal region of a body includes an ultrasound linear array configured to generate ultrasound pulses, to receive and process ultrasound reflection data, and to receive and process photoacoustic signals; a laser configured to generate narrow-beam laser pulses having a pulse energy of less than 5 mJ; a laser pulse distributor configured to receive laser pulses from the laser and to direct the pulses to a plurality of locations on a target; and a computer system configured to control the ultrasound linear array, the laser, and the laser pulse distributor, to receive processed ultrasound reflection data and photoacoustic signals from the ultrasound linear array, and to generate combined photoacoustic and ultrasound images from the received data and signals.

DETAILED DESCRIPTION

A particular embodiment of an integrated scanning photoacoustic and ultrasound imaging system and method will now be described with reference to the drawings, wherein like numbers indicate like parts. In the embodiment shown inFIG. 1, a low-power, small-diameter beam, pulsed laser is scanned over a region of interest, and a photoacoustic sub-frame (or sub-image) response is recorded for each laser pulse, as discussed in more detail below. Between laser pulses several sequential ultrasound pulse-echo beams are formed close to the scanning laser beam location after each laser pulse, and pulse-echo sub-frame images are recorded.

There are two benefits of interleaving ultrasound beams between photoacoustic sub-frames: 1) the ultrasound information and the photoacoustic information are acquired at the same or nearly the same spatial location; this is especially important for imaging moving objects or monitoring fast changing dynamics; and 2) the dead-time between laser pulses (tens of microseconds photoacoustic recording for centimeters range versus the millisecond PRI for a kilohertz laser) is fully utilized, increasing the frame rate because ultrasound pulse-echo recordings do not have to wait until all photoacoustic recordings are finished. This novel method of interleaving ultrasound sub-frames between photoacoustic sub-frames enables the production of real-time interleaved photoacoustic/ultrasound images for deep imaging (up to several centimeters) with a frame rate greater than 30 Hz, thereby enabling the use of such combined images for clinical applications. Frame rates of 60 Hz or greater may be readily achieved in real-time to produce moving images of an internal region of interest.

For example, an integrated photoacoustic/ultrasound imaging system disclosed herein is suitable for clinical applications where co-registered photoacoustic/ultrasound is needed to provide real-time and precise monitoring with high-contrast images. As discussed below, the system disclosed herein has produced real-time video showing needle insertion into tissue and simulated therapeutic agent injection.

Ultrasound (only) image-guided needle monitoring has been used in clinical applications such as biopsy in prostate and breast, and therapeutic agent injection. Unfortunately, the quality of the ultrasound image is directly related to the orientation of the needle relative to the scan direction of the real-time image because the primary signal from a needle is the specular reflection from its surface. In many cases, proper orientation is problematic and tracking precision is lost because of the poor ultrasound contrast and additional artifacts associated with the needle. In addition, background scattering degrades image contrast, limiting the precision of the biopsy and injection.

However, photoacoustic signals from the needle are much less dependent on the orientation of the light source and less sensitive to needle-transducer orientation. In photoacoustic imaging the needle is a cylindrical acoustic source in the transducer imaging field. In ultrasound imaging, however, the needle is a reflector and usually scatters the probe ultrasound beam in a direction different from that of the incident ultrasound beam. As a result, photoacoustic imaging is less sensitive to transducer-target orientation compared with ultrasound imaging. Also, photoacoustic imaging exhibits higher contrast to monitor injected exogenous agents because of the speckle-free background, and the ability to easily label agents with molecular dyes of high optical absorption at the optical wavelength used for photoacoustic imaging.

An experimental arrangement demonstrating image acquisition for an interleaved photoacoustic/ultrasound imaging system100in accordance with the present invention is shown schematically inFIG. 1. A block diagram of the system100is shown inFIG. 2.

The system100includes an ultrasound probe102having a linear array with 128 channels (e.g., AT8L12-5 50 mm, Broadsound Corp.; central frequency 9 MHz, bandwidth 5 to 12 MHz, 256 elements, 195 μm pitch, elevation focus 20 mm). The linear array is used for both ultrasound pulse-echo imaging and photoacoustic signal reception. A diode-pumped laser104(e.g., TECH-1053 Specific, Laser-Export Co. Ltd.) delivers 10-ns pulses with a pulse energy of 1.9 mJ at a wavelength of 1053 nm with a 2.8-mm beam diameter, resulting in a fluence of 31 mJ/cm2at the sample surface. This fluence is well below safety limits, 100 mJ/cm2, at this wavelength. It will be appreciated by persons of skill in the art that suitable laser systems have the advantage of compact size and low cost. The maximum pulse repetition rate in this embodiment is 100 kHz, with maximum output pulse energy of about 2.2 mJ at 1 kHz.

The beam from the laser104is directed to a galvanometer scanner106, referred to herein as a galvo-mirror (e.g., GVS001, Thorlabs Inc.). In this embodiment the mirror angle for the scanner106is varied between −1° and +1° using a sinusoidal waveform delivered by a function generator122(e.g., AFG 3252, Tektronix Inc.). At a distance of 30 cm from the target this angular range yields a lateral scanning range of about 1 cm. In another embodiment the galvo-mirror angle is varied using a triangle waveform, to produce a more uniform coverage of the image region.

Two constraints when choosing the illumination condition are: 1) the laser fluence should be below safety-related wavelength-dependent thresholds (e.g., 100 mJ/cm2at 1053 nm wavelength) for a single laser pulse, and 2) the average power should be under 1 W/cm2at every sample point. For example, to meet real-time frame rates (e.g., 30 Hz or greater), the fluence is limited to 33 mJ/cm2(1 W/cm2/30 Hz) because the laser spot returns to the same position every 1/30 seconds. A tradeoff between SNR (related to fluence) and frame rate was considered in designing the illumination sequence.

In this exemplary system a 30-Hz frame rate (mirror scan rate) is set, with 24 scanning beams105over a 1 cm range and a laser pulse repetition rate of 720 Hz. The axial imaging range is set to 2.8 cm. Both the laser104and the ultrasound probe102are operatively above the test phantom90and aligned such that the incident plane of the laser light105and the ultrasound probe imaging plane103intersect at a target depth107. In this embodiment the laser104is configured such that the beams105are tilted by about 45° with respect to the ultrasound probe102.

The clinical applicability of this system100is demonstrated for needle guidance by providing substantially real-time imaging of needle96insertion into a test phantom comprising a chicken breast92(approximately 3 cm in thickness). The chicken breast92was surrounded by a 10% acoustically and optically transparent gelatin shell94(1 cm thickness) for ultrasound coupling. An 18-gauge needle96(e.g., 1.27-mm outer diameter, 18G1, Becton Dickinson & Co., East Rutherford, N.J.) is mounted on a linear stage (not shown) for insertion laterally into the test phantom90with a tilt angle of about 20°. The target depth in this embodiment is up to 12 mm inside the biological tissue (i.e., chicken breast92). With a laser light-tissue incident angle of 45°, the light propagation path into the chicken breast92is about 17 mm. To demonstrate the real-time imaging capability of the system100, the insertion of the needle96and the injection through the inserted needle96of an ink solution (e.g., 44011, Higgins Ink, Leeds, Mass.) are imaged, simulating delivery of a therapeutic agent such as a small molecule drug, combined with a molecularly absorbing dye. The ink solution had an optical absorption coefficient of 20 cm−1at 1053 nm.

A block diagram of the system is shown inFIG. 2. A programmable ultrasound scanner110(e.g., Vantage, Verasonics Inc., Redmond, Wash.) is used for all image data acquisition and reconstruction (photoacoustic and ultrasound). The system100is operated using Matlab® (The MathWorks Inc., Natick, Mass.) on a conventional computer120. The computer120controls ultrasound and photoacoustic transmits, receives, and data transfer events. A function generator122is used to synchronize all operations (i.e., laser pulses, photoacoustic signal recording, and mirror position). In addition to sending a control function signal (e.g., a sine wave signal) to drive the galvo-mirror106, it sends triggers to the laser104and the ultrasound scanner110simultaneously to synchronize photoacoustic signal recording with each laser pulse. During one period of the sine wave, multiple laser pulses (i.e., multiple trigger signals) are emitted, covering the entire lateral imaging range (e.g., approx. 1 cm). After each laser pulse and photoacoustic signal recording, several ultrasound pulse-echo beams are formed. The transmit ultrasound pulse selected for this embodiment has a center frequency of 7.8 MHz, and the sampling frequency on receive is 31.25 MHz with a bandwidth of twice the center frequency.

To optimize the frame rate, a large data set (corresponding to several mirror cycles) is acquired and then transferred at once, while the next data set is being acquired, providing continuous data acquisition, as discussed in more detail below.

The particular parameters and exemplary components identified herein are disclosed to provide a detailed disclosure of a particular embodiment to aid persons of skill in the art in understanding the invention. Other parameters and similar components will be apparent to such persons of skill in the art, and the present invention is not limited to these particulars.

An exemplary sequence of steps for scanning an internal region of interest150within a body152in accordance with the present invention is shown inFIGS. 3A-3F. In this example the internal region of interest150is disposed at a depth D from a surface154of the body152. For example, the depth D may be in the range of 1-6 cm from the outer surface154. In another example, the depth D may be about 2-4 cm from the outer surface154of the body152.

A first location154A on the outer surface154is irradiated with a small-diameter, low-pulse-power laser pulse105A that is configured to irradiate at least a portion of the internal region of interest150, inducing a photoacoustic response155A (FIG. 3A). Although the pulse-power of the laser154A is low, the laser beam is sufficiently narrow at the outer surface154that a high-SNR can be obtained. However, it will be appreciated that due to high optical scattering in tissue a significant portion of the region of interest150is irradiated. A plurality of ultrasound pulse-echo beams156A are then sequentially directed towards the outer surface154and region of interest150, and corresponding ultrasound reflection signals157A of the ultrasound beams156A are obtained (FIG. 3B).

A second location154B on the outer surface154is then irradiated with the similar laser pulse105B, inducing another photoacoustic response155B (FIG. 3C). A second plurality of ultrasound beams156B are then sequentially directed towards the outer surface154and region of interest150, and corresponding ultrasound reflections157A are obtained (FIG. 3D).

The process is repeated during the scan (FIGS. 3E and 3F), wherein a predetermined nth laser pulse154N irradiates an nth location154N on the outer surface154, producing a photoacoustic response155N. A series of ultrasound beams156N are subsequently directed sequentially to the outer surface154, and ultrasound reflections157N are obtained, completing one-half of a mirror cycle.

The photoacoustic signals for each cycle are combined to produce photoacoustic sub-frames, and the corresponding pulse-beam ultrasound signals are combined to produce ultrasound sub-frames. The photoacoustic sub-frames and the ultrasound sub-frames are combined to produce a substantially real-time image of the region of interest that combines the advantages of photoacoustic imaging with the advantages of ultrasound imaging.

An exemplary detailed timing diagram200of the acquisition for one data set is shown inFIG. 4, wherein the system100runs continuously by repeating this acquisition set. The timing diagram200illustrates the mirror position202for one acquisition set comprising a plurality of mirror cycles203that scan the pulsed laser beams105over the region of interest. One mirror cycle203in this example comprises 24 laser pulses or pulses204, followed immediately by photoacoustic signal acquisition. Each laser pulse204(and photoacoustic signal recording) is followed by a plurality of (six in this example) ultrasound pulse-echo ultrasound beams (labelled1-6,7-12,13-18,139-144) inFIG. 4. Therefore, one mirror cycle203produces one photoacoustic frame with 24 sub-frames, and one ultrasound frame from 144 sub-frames. One mirror cycle203is completed in 33.3 ms, resulting in an integrated frame rate of 30 Hz.

The number of mirror cycles constituting one data set is selectable and is limited by the maximum amount of data that can be transferred at once by the hardware of the ultrasound scanner110. For each mirror cycle203, there are 24 laser pulses for photoacoustic receives (128 channels for each laser pulse). Each laser pulse/photoacoustic receive is followed by six full beams of pulse-echo ultrasound focused at six different adjacent lateral locations. Laser pulses are preceded by adding a flush time107(no operation, e.g., 100 μs in current embodiment) to prevent ultrasound emissions from interfering with photoacoustic signal recording. This pattern is repeated until one mirror cycle is completed, allowing the laser beam to cover the entire lateral imaging range twice, and leads to: (i) one integrated photoacoustic frame, resulting from a total of 24 laser pulses at a 720 Hz repetition rate (i.e., 24 single shot sub-frames); and (ii) one integrated ultrasound frame consisting of 144 ultrasound pulse-echo focused beams swept laterally across the imaging range, similar to conventional B-mode ultrasound imaging.

In this embodiment the ultrasound scanner110data transfer time from the front end to the host computer120memory is determined by a transfer rate of 6600 MB/s and also includes a few milliseconds (typically 5 ms) overhead for each transfer. To reduce the number of transfers, a large data set is acquired and transferred together. The data sets are acquired continuously; while one data set is transferred, the next one is acquired. Parameters are selected such that the data transfer time is shorter than the acquisition time, to allow for continuous data acquisition and real-time imaging.

In such a continuous mode, the frame rate is governed by the galvo-mirror106period. One mirror cycle lasts 24/720 Hz=33.33 ms, which corresponds to a frame rate of 30 Hz for an integrated photoacoustic/ultrasound frame for a field of view of 2.8×2.8 cm (lateral×axial). The laser104scan range does not cover the whole photoacoustic image range. Nevertheless, light scattering in tissue allows for almost homogeneous irradiation at a few millimeters depth and may cover the entire image range at the imaging depth (e.g., ˜1 cm or greater). Both photoacoustic and ultrasound images were reconstructed using a delay-and-sum beam forming approach, Hilbert transformed, with the envelope displayed on a logarithm scale. A fusion image was also obtained by overlaying a photoacoustic image on the top of the corresponding ultrasound image (i.e., add photoacoustic pixel value to corresponding ultrasound pixel and display with different color mapping).

Detailed results from the system100and method discussed above are provided in C.-W. Wei, et al., “Real-Time Integrated Photoacoustic and Ultrasound (PAUS) Imaging System to Guide Interventional Procedures: Ex Vivo Study”, IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, Vol. 62, No. 2, February 2015, which is hereby incorporated by reference.

The results illustrate the system100and method provide real-time images showing the needle96entering the tissue92, retracting within the tissue92, and injecting the ink marker into the tissue92.

With ultrasound imaging alone the image of the needle96is only slightly brighter than the tissue92, with limited contrast because of background scattering. The needle96is barely distinct from the trace in static ultrasound images. In contrast to ultrasound, the integrated photoacoustic image exhibits more than 30 dB contrast over the background. Two disconnected segments of the needle are visible in the integrated image and correspond to the geometry and structure of the needle96, with a tilted injection hole at the needle96front. The top segment corresponds to the top edge of the needle96, and the lower and shorter segment is the lower tip under the hole, with the proximal part not seen because light was blocked by the top needle edge. The rim of the hole has a weaker signal because it was away from the image plane in the elevational direction.

To fully mimic an interventional procedure using a needle, injection of an absorptive agent was imaged ex vivo. A movie showing the needle96entering a piece of chicken breast tissue92, injecting ink solution, and exiting from the tissue92was generated, with four sections, corresponding to: (i) inserting the needle96into the phantom90from a top right corner at an angle of about 20° towards the lower left corner, to a maximum depth of about 22 mm (12 mm below the tissue surface); (ii) retracting the needle96until the tip is at a depth of about 18 mm; (iii) injecting the ink solution; and (iv) removing the needle96from the tissue92.

Right after injection, the needle tip96is clearly seen in the combined photoacoustic image/ultrasound image. Reverberation artifacts (i.e., multiple reflections) are also observed Ink pours out and diffuses to the volume surrounding the tip, and then flows through the channel in the tissue92created when the needle96was inserted. At the end of the injection, the diffused ink masks the needle tip96. The contrast of the needle/injected ink to the tissue92background on the combined photoacoustic/ultrasound image is more than 30 dB with a laser penetration distance of more than 1 cm, demonstrating the feasibility of real-time photoacoustic/ultrasound image-guided drug delivery.

A wavelength of 1053 nm was chosen in this study because it provides good penetration and high contrast among commonly used wavelengths (700 to 1100 nm) in photoacoustic imaging. For needle guidance, the optical wavelength is not critical because of the strong optical absorption in metals over the whole therapeutic range of wavelengths. In addition, inexpensive, compact, high powered, fast commercial fiber or diode-pumped lasers operating at this wavelength are available, making clinical translation easier. Wavelength-tunable laser systems enabling spectroscopic imaging, such as the Opolette HR 532 (Opotek Inc., Carlsbad, Calif.) with length >60 cm including power unit+laser head, and versaScan OPO (Newport Corp., Irvine, Calif.) with length ˜30 cm, are also available, and the scanning approach can be easily implemented using these laser systems. Compared with these laser systems, usually requiring an additional pumping laser, however, a single-wavelength system is still more compact. For example, exogenous absorbers contain a magnetic component and thus can be magnetically manipulated. By detecting the magnetically induced changes (e.g., displacement), exogenous absorbers can be differentiated from endogenous ones, which are insensitive to the magnetic field.

Although single-wavelength imaging in the region around 1 μm will have many applications, a robust photoacoustic molecular imaging system will benefit from wavelength tunability. In another embodiment the laser110is a wavelength-tunable laser, as are known in the art. Leveraging the same laser diode-pumped technology exploited here to serve as the pump for an optical parametric oscillator is contemplated for a system operating at high pulse repetition rates.

In the current embodiment the laser beam104angle is at a relatively large angle (approximately 45 degrees) with respect to the ultrasound imaging plane defined by the ultrasound scanner110. In another embodiment a fiber-optic system250is combined with the ultrasound probe102in an integrated photoacoustic/ultrasound probe252, as shown inFIG. 5. The laser pulse104is distributed to a fiber bundle254that transmits the laser pulse along either, or preferably both, sides of a conventional ultrasound array (e.g., on both sides of the elevation aperture) to optimally illuminate the image plane. The narrow beam laser pulse from the high repetition-rate laser104is distributed sequentially to different fibers256in the bundle154to sweep the small-diameter laser beam over the surface of the body. The sweep is achieved, in this exemplary embodiment, in a remote switching unit260in which the laser104output is injected to a different fiber256on each laser pulse.

The disclosed narrow-beam scanning approach provides image quality that is substantially equivalent to the image quality achievable using a broad-beam irradiation for the same illumination area. The photoacoustic images are summed coherently, i.e., raw wave field after delay-and-sum beam forming but before Hilbert transformation. Because delay-and-sum is a linear operation, summing the wave field for each narrow laser beam is equivalent to the reconstructed wave field generated by a broad beam summed with all the narrow beams (Huygens-Fresnel principle). For imaging a fast-moving object, motion artifacts may be present. Nevertheless, no significant artifact should be observed because light is highly scattered at the imaging depth (˜1 cm) and the illumination patterns from multiple beams overlap over the 1-cm scan range.

The frame rate is currently limited by the laser average power regulation, 1 W/cm2, as mentioned above. To reach a frame rate larger than 30 Hz, the laser fluence must be smaller than 33.33 mJ/cm2. For a higher frame rate, the fluence decreases and thus image SNR degrades. Nevertheless, with a high SNR (>35 dB for needle guidance), it is possible to increase the frame rate at the price of lower laser fluence (lower SNR), e.g., 50 Hz frame rate and 20 mJ/cm2fluence. Also, in the current setup, the mirror scans from one end to the other and returns within one cycle, so two full photoacoustic frames can be formed, instead of one. Therefore the frame rate may be doubled, albeit with the SNR degraded by 3 dB (½). An alternate way is to scan in two dimensions with an additional galvo mirror to scan elevationally, in addition to lateral scanning, i.e., zigzag scan. The laser beam will not return to the same spot at each lateral scan cycle. The fluence can be kept constant while the laser PRF is increased and, thus, the frame rate increased.

The unique mechanism of optical absorption-based contrast makes photoacoustic imaging an ideal modality for molecular imaging, while ultrasound can show simultaneous anatomical landmarks, providing context for molecular measurements. A particular example is to detect circulating tumor cells in blood vessels, providing a critical indicator of metastasis for cancer. Metastatic tumor cells circulating in the vasculature can be accumulated and detected using a composite nanosystem exhibiting both strong magnetic properties and high optical absorption at a desired wavelength. By targeting these particles to specific biomarkers on tumor cells, circulating tumor cells can be trapped and manipulated with an external magnetic system and non-invasively imaged using the integrated photoacoustic/ultrasound imaging disclosed herein. The imaging site is chosen at the radial artery in the forearm, which has a diameter of 1 to 2 mm and a level of 10 mL/min flow rate and, thus, a reasonable examination time to interrogate a significant blood volume. The accumulated cells create heterogeneous photoacoustic sources inside vessels, and thus can be detected by a commercial ultrasound probe with a narrow frequency band. In short, the system presented in this paper is well-suited to clinical translation of this technology.