Patent Application: US-201213353797-A

Abstract:
a method and system for inspecting biological tissue that has no applied coatings or treatments to improve reflectivity comprises an optical detection system with an exposed surface for inspection by an optical detection system ; and a laser for exciting an ultrasonic wave within the tissue , which wave propagates within the tissue at least near the surface . the optical detection system includes : a laser to emit a pulsed detection laser beam onto the surface at a detection spot , the pulsed laser beam having a wavelength at which there is large scattering and little absorption by the tissue , and a pulse duration is chosen to correspond with ultrasonic propagation times associated with a range of depths of the scan , whereby the fluence of the detection laser is not applied unless ultrasonic information regarding a given depth is being obtained ; and large étendue collection optics for collecting reflected and backscattered light from the detection spot ; and a demodulator to extract information from the ultrasonic wave from the collected light .

Description:
a technique is described for non - contact detection of ultrasonic waves propagating in biological tissues . a schematic diagram of a preferred embodiment is shown in fig3 . a biological tissue ( bt ) of human or animal origin , either in vivo or ex vivo , has an unprepared surface . the bt is immobilized with respect to the examination equipment in any of the various ways known in the art , to stabilize the bt . immobilization is required only for the duration of the scan , and therefore if the scan is quick , and the tissue is structurally supported , then immobilization may be effectively provided by supporting the equipment with a table in the same room as the tissue . the tissue immobilization may be provided by placing the tissue in a dish or plate , and preventing it from moving during the scan . the scan may be a very short a scan , in which case the scan may be expected to work even with some motion , and a hand - held probe can be used . longer b scans may require some equipment for mapping the distances to present the scan information in a registered manner , but still may be provided by hand - held probe . c scans may require more registration but may not require any clamping or contact with the tissue . for surgical interventions , bodies are generally immobilized to a much higher degree than required for imaging or inspection according to the present invention . for endoscopic applications , sufficient immobilization is generally provided with known means . a short pulse ( typically 5 - 20 ns , but for some tissues , laser wavelengths and glb power , it may range from 1 - 50 ns ) generation laser beam ( glb ) illuminates the surface on a generation area , which is typically of a largest size available for the tissue to be imaged , so a diameter of about 2 cm for small mammals , or larger areas for larger mammals . the wavelength of the glb is chosen to be within the optical window of biological tissues that is between about 500 nm and 1000 nm in order to increase the probing depth of the photons . due to the bt composition ( cells and other optical scatterers ), the glb is strongly scattered . most photons are backscattered just below the surface and are lost in air after only a few scattering events . the remaining photons penetrate more deeply into the bt , following highly randomized optical paths . photons of the glb will thus produce ultrasonic waves in two different ways . firstly , any residual absorption of glb photons backscattered in the immediate vicinity below the surface , where the power density is much higher , will generate an ultrasonic wave by thermoelastic generation . this produces an ultrasonic wave ( uw ) with a wavefront similar to the surface topography having a lateral extent given by the glb diameter . after propagation , this uw will be scattered by any acoustic inhomogeneity within the bt , thus leading to backscattered ultrasonic waves . secondly , glb photons diffusely propagating more deeply inside the bt will also produce ultrasonic waves in presence of any optical absorber . those ultrasonic waves then propagate throughout the bt in all directions until they reach the unprepared surface . the surface is monitored with a pulsed detection laser beam ( dlb ) which is generally focused on a smaller area than the generation area ( typically 1 / 10 th to 1 / 100 th or more ) in order to provide a spatially resolved temporal measurement of the surface displacement at discrete locations . these measurements are then used to generate a 2d or 3d image characterizing the tissue by using mathematical procedures known in the art . the dlb has a wavelength larger than 1 μm due to the higher tolerance of bt to laser exposure at wavelength above 1 μm , and particularly a wavelength from 1 - 1 . 8 μm , and more preferably 1 . 5 - 1 . 8 μm ( ref . 10 ). the wavelength of 1 . 06 μm is often preferred due to the availability of suitable laser sources already developed for previous - art industrial and metrology applications . the biological nature of the tissue restricts the amount of laser energy which can be used for both the glb and the dlb , however because the glb is not constrained by the resolution of the image , the dlb fluence is critical . the pulse duration of the dlb is thus chosen to reduce the tissue exposure to laser light . depending on an intended scan depth , or range of depths to be scanned , a corresponding pulse duration is required . the pulse duration corresponding to a range of depths may be triggered at a time delay correspondence to the ultrasonic propagation time ( e . g . ˜ 1 . 5 mm / μs ), from the surface to the depths , in a manner well known in the art . reducing the pulse duration allows increasing the peak power incident on the medium , thus increasing the sensitivity of the measurement . typically , the dlb pulse duration will be less than 20 μs , corresponding to a propagation distances between 15 to 30 mm in typical bt . for applications in which the depth range of interest is much shorter , e . g . probing vascularisation within eye retina with detection directly off the retina surface , off adjacent layers within the eye or even from the surface of the blood vessel , the dlb pulse duration could be as short as 1 μs or shorter . dlb light which is reflected / backscattered by the bt is collected using a large étendue optical system and processed by embodiments known in the art . the large étendue collecting optical system may be coupled to a multimode optical fiber of comparable étendue . typically , an optical fiber with a core diameter of 1 mm or below and a numerical aperture ( na ) below 0 . 4 will be sufficient . this corresponds to a maximum étendue u max =( πφna / 2 ) 2 = 0 . 4 mm 2 sr , where φ is a core diameter of the optical fiber . such a value is comparable to the étendue of known optical demodulators such as a cfpi ( 50 to 100 cm long cavity ) or a pri . phase demodulation with spectral hole burning in cryogenically cooled rare - earth ions doped crystal would also be suitable but more complex to implement , especially because of the difficulty of matching the atomic filter wavelength to that of the dlb . the mathematical processing related to backscattered ultrasonic waves or pat - like emitted waves are both known in the art . the two imaging modalities may be both employed for analysis of a same signal . both imaging modalities are schematically illustrated in fig4 a ), b ). firstly , imaging in ncpat mode , like in conventional pat , will provide information on the location and the extent of optical absorbers ( inclusions ) embedded in the biological tissue , by generating an opto - acoustic wave . in this case , schematically shown in fig4 a ), the ultrasonic wave is coming from the inclusion to the surface , and corresponding saft reconstruction relies on the one - way time of flight between the optical absorber and the surface ( dotted line in fig4 a )). secondly , imaging in ncus mode will provide information on the location and extent of acoustic discontinuities ( inclusions — though may be entirely different from the absorbers of fig4 a )) within the bt . in this case , schematically shown in fig4 b ), the time of flight in the saft algorithm is the sum of the time of flight of the ultrasonic wave from the surface to the discontinuity and that of the backscattered ultrasonic wave from the discontinuity to the surface ( dotted line in fig4 b )). for both imaging modalities , the topography of the surface must be taken into account for a better reconstruction . in most real applications , embedded objects or inhomogeneities will present both acoustic impedance mismatch and optical absorption contrast with the surrounding medium . the use of both image reconstruction algorithms will thus provide two types of information on each embedded object . the choice of the glb wavelength as well as the intrinsic properties of the bt will also influence the relative strength of the backscattered ultrasonic wave and the opto - acoustic ultrasonic wave . fig5 schematically illustrates two envisioned systems which could be easily derived from the preferred embodiment for applications for in situ diagnostics on unprepared surfaces of biological tissues , during examinations or surgical procedures as well as for small animal diagnostics performed in research laboratories . fig5 a ) shows a drawing of a handheld probe linked to the main laser system by an articulated optical arm including folding mirror and relay lenses , or an optical fibre to provide a continuous waveguide at any state of extension . the user would scan the surface , either manually ( with a tracking system based on position encoders or by image tracking of the detection spot with a ccd camera ) or with a scanning device inside the handheld probe . from the tracked position an image ( b - scan or c - scan ) can then be constructed . the second implementation shown in fig5 b ) uses a fiber optic bundle ended by micro - optical components in order to perform ncus and ncpat diagnostics on parts of the body accessible through the mouth , the respiratory airway , the gastro - intestinal track , or other orifices , or through small incisions in the skin for less invasive surgical procedures . the application to retinal diagnostics may involve directing the glb and dlb essentially unscattered through the cornea to the retina . a signal related to the absorption of the glb by blood in a vessel is obtained by detecting the reflection or scattered light from the dlb off the interface retina - vitreous humour or the surface of the vessel itself or any interface between the various layers at the back of the retina ( retinal pigmented epithelium , choroid and sclera ). detecting directly off the surface , the surface motion of the vessel at ultrasonic frequencies , requires the shortest duration of the dlb , since this motion may occur very quickly . the duration of the surface motion of the vessel varies with the size of the vessel and the penetration of the glb in blood . its duration is calculated to be below 50 ns when the glb is at 532 nm and therefore , a suitable duration of the dlb of 1 μs and even shorter is preferred . this short duration allows for the use of higher peak power , which translates into higher sensitivity , as mentioned above . calculations we have performed , which are similar to one indicated above , have shown that probing blood oxygenation with non - contact detection off internal layers of the eye is feasible and could be performed safely . these calculations , the experimental results obtained so far , and ocular safety limits ( see [ 10 ]) support the conclusion that ncpat and ncus can be used for retinal diagnostics . such an approach advantageously replaces detection with an ultrasonic transducer contacting externally the eye , as reported in [ 15 ]. fig6 is a schematic illustration of a system used to verify the present invention . the generation laser was a frequency doubled actively q - switched laser emitting 5 ns pulses at 532 nm . this wavelength efficiently excites blood , making it an effective ultrasonic transducer by thermoelastic generation . the repetition rate was equal to 10 hz and the pulse energy impinging on the surface was about 100 mj . the glb was oriented toward a bt onto a large spot size ( about 25 mm in diameter giving an area of 5 cm 2 ) which satisfies the maximum permissible exposure ( mpe ) at generation wavelength ( 20 mj / cm 2 at 532 nm ). with a 10 hz repetition rate , the average power density of the glb was equal to 200 mw / cm 2 , which is also in accordance with the safety limit in the case of repetitive exposure at 532 nm . the detection laser beam was transmitted by a second optical fiber and focused onto the surface of the tissue after reflection by a computer controlled scanning mirror . the dlb had a wavelength λ of 1 . 06 μm . the pulse duration was set according to the maximum ultrasound propagation time for the desired probing depth and the peak power and the laser power was set to approach the safety limit as described above . scattered light was collected by a confocal fabry - perot interferometer ( cfpi ) used in transmission mode . the transmission mode was chosen for its higher sensitivity at low frequencies ( below 10 mhz ). a 1 meter long cfpi was used with mirror reflectivities r = 94 . 5 %, giving a finesse f = πr /( 1 − r 2 )= 28 and an optical étendue estimated as u = π 2 lλ / f = 0 . 375 mm 2 sr , where l , the length of the cfpi cavity , is 1 m . the experiments used two collecting optical fibres . the first had a core diameter φ = 1 mm and a numerical aperture na = 0 . 39 , giving an étendue u =( πφna / 2 ) 2 = 0 . 375 mm 2 sr comparable to that of the cfpi . the second optical fibre , used in other experiments , had a na = 0 . 36 and φ = 0 . 4 mm , thus giving u = 0 . 051 mm 2 sr . as shown in fig6 , the coupling of the dlb in front of the collection optics ensures good separation between illumination and collection , and avoids stray light from illumination being coupled into the detection channel . the inset in fig6 shows schematically a top view of the sample , of the large size generation area and of the scanned small detection spots . although the embodiment used a mirror scanning the detection spot , it also possible to have a configuration in which a larger detection spot is not scanned . as shown in fig7 , the detection laser used in our implementation was composed of a master oscillator ( mo ) which was a commercial continuous wave single - frequency nd : yag laser emitting about 200 mw at 1064 nm followed by a multipass flashlamp pumped nd : yag amplifier . an intensity modulator ( imod : electro - optic modulator in our implementation ) was located between the mo and the amplifier . the electrical signal feeding the imod was tailored to produce a top hat temporal profile of appropriate duration . the amplifier was a three - pass configuration using two nd : yag laser rods ( lr 1 and lr 2 located in a pumping chamber including a flashlamp ( fl ) seeded by a variable pulse length power source ( not shown ). two optical isolators were used : the first one ( oi 1 ) was located between the imod and the amplifier to protect the mo from any unwanted optical feedback from the amplifier , the second one ( oi 2 ) was located before the third stage of amplification to eliminate self oscillation of the amplifier and the accompanying relaxation oscillations . the quarter - wave plate ( qw ) and the thin - film polarizer were used in combination to use the lr 1 for the first two amplifying stages . short pulse duration ( shown in fig8 ) is obtained by using the electro - optic imod between the master oscillator and the optical amplifier . alternatively the switch could be introduced at the output of the amplifier , although having it located before amplification is generally more efficient for extracting the energy stored in the amplifier while tailoring the temporal shape of the output pulse . alternatively also , pulse shortening can be obtained by limiting the duration of the electrical pulse feeding the pumping flashlamp ( s ) or the pumping laser diode ( s ) in the single or multistage optical amplifier . in such a case , however , the pulse shape is primarily determined by the pumping flash temporal profile and the amplifying medium dynamics ( laser transition lifetime ). tailoring the temporal profile of the detection laser pulse can also be used to compensate for tissue ultrasonic absorption and limited glb penetration by having power increasing with time . this is equivalent to time - gain compensation ( tgc ) used in conventional ultrasound . the detection laser pulse energy is thus used optimally to obtain the best sensitivity while minimizing the exposure to laser light . with each laser shot , an a - scan signal is obtained . the scanning mirror ( one - axis or two - axis ) allows scanning over the surface and then from the a - scan signals , b - scans and c - scans can be plotted as in conventional us . reconstruction techniques such as saft and back projection algorithms are used to get higher resolution , two - dimensional ( 2d ) or three dimensional ( 3d ) mapping of embedded inclusions or anomalies within the tissue . the embodiment described above without the scanning mirror was used to demonstrate the ncus and ncpat diagnostic modes on a phantom having similar optical properties as human or mouse tissue . this phantom was made of pieces of raw chicken breast with inclusions such as a steel plate , black polyethylene ( pe ) strip or black jelly embedded from 10 to 20 mm deep below the surface . fig9 shows the results obtained for the ncus mode in which the phantom surface was slightly tinted with red dye to simulate the presence of blood at its surface . an echo coming from the steel plate is easily visible at time delay of approximately 22 μs , which corresponds to a propagation distance of about 17 mm . the size of the detection spot was about 3 mm in diameter . the same system was also used in ncpat mode . the inclusions in the phantom were a piece of black jelly of 25 × 3 × 3 mm 3 and a thin sheet of black polyethylene strip located respectively at about 23 mm and 12 mm below the surface . in this case , no red dye was used on the surface of the tissue , thus allowing the generation laser pulse to illuminate the whole volume of the sample . the demodulated signal is shown in fig1 . a first peak is seen at a time delay of about 8 μs , corresponding to the depth of the black strip . the second peak , which is wider , is located at a time delay of about 17 μs , also corresponding to the known depth of the black jelly . in this case , the duration of the peak is much longer since every point of the long jelly sample acts as an ultrasonic wave source . consequently , points at both extremities contribute to the signal at long delays whereas the point at the center corresponds to the shortest propagation delay . the other examples show the capability of the approach to produce good resolution b - scan images by scanning a smaller spot and using reconstruction by saft processing . to generate this data , scanning was performed by moving the phantom instead of using a scanning mirror . fig1 shows the preparation of a chicken breast phantom with 3 embedded black polyethylene strips as inclusions . these inclusions were covered by an additional layer of chicken breast tissue ( about 12 mm thick ). fig1 shows the raw b - scan data ( upper image ) obtained on this phantom and the saft processed data for ncpat mode ( lower image ). after reconstruction , the 3 absorbing inclusions are clearly seen . in this case , the detection spot size was 1 mm in diameter and the scanning step was 1 mm . 41 a - scans were used for the reconstruction . fig1 shows results obtained on a chicken breast phantom with a single 2 mm wide black polyethylene strip embedded about 15 mm below the surface . they include the raw b - scan data ( upper image ), the saft processed b - scan in ncpat imaging mode ( middle image ) and ncus imaging mode ( lower image ). when ncpat is assumed and saft is applied accordingly , the inclusion is clearly seen after processing although not distinguishable in the raw data . when a source of ultrasonic wave is assumed to be close to the surface , processing for the ncus imaging mode shows clearly a small impedance mismatch between the two pieces of chicken breast . in this case the surface has not been dyed but the chicken breast has however some background absorption at the 532 nm wavelength used for generation . since there is a strong increase of light energy density close to the surface due to scattering , it is then reasonable to assume that there is a source of ultrasound near the surface . in this case interrogation of the phantom is purely acoustic and does not indicate a buried optically absorbing inclusion . these data were obtained with a detection spot diameter and scanning step of 0 . 4 mm . 81 a - 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[ 15 ] h . f . zhang , c . a . puliafito , s . j . jiao , “ photoacoustic ophthalmoscopy for in vivo retinal imaging : current status and prospects ”, ophtahalmic surgery , lasers & amp ; imaging , vol . 42 , pp . s 106 - s 115 ( 2011 ). other advantages that are inherent to the structure are obvious to one skilled in the art . the embodiments are described herein illustratively and are not meant to limit the scope of the invention as claimed . variations of the foregoing embodiments will be evident to a person of ordinary skill and are intended by the inventor to be encompassed by the following claims .