CN116482853A - Light imaging and processing device for constructing thin light layer in turbid medium by multi-beam interference - Google Patents

Abstract

The present invention relates to a device for optically imaging or processing a target in a turbid medium. The invention utilizes multi-beam interference to generate coherent destructive interference in the propagation paths of illumination, processing and signal beams to reduce the combined intensity of the illumination, processing and signal beams, thereby reducing the light absorption and light scattering of turbid media; and at a target position in the turbid medium, the synthesized light intensity of multiple light beams generated by utilizing the coherent constructive interference of the multiple light beams is extremely high, so that an internal thin light layer is formed to illuminate and process the target. Theoretical analysis and mathematical calculation show that the device has excellent performance. The imaging or incision-free surgical depth of the device in the human body can exceed 5 cm, and the imaging or wireless optical communication distance in clear seawater can exceed 500 m. The imaging resolution of the device along the object plane during imaging can be equal to or even smaller than the wavelength of visible light, and the imaging resolution in the depth of field direction of the target can be close to the wavelength of visible light.

Description

Light imaging and processing device for constructing thin light layer in turbid medium by multi-beam interference

Technical Field

The present invention relates to an apparatus for optically imaging and processing objects in turbid media, and more particularly to an apparatus for optically imaging and processing objects by illuminating or processing objects in turbid media by means of multi-beam interferometry by constructing an internal thin light layer and a method based on the same.

Background

Imaging in turbid media, including optical imaging, acoustic imaging and X-ray imaging, has found widespread and urgent need in modern society, such as medical imaging, underwater imaging, imaging in haze, and detection of internal injuries to solid materials, etc.

Also, since laser processing can modify, vaporize, ablate, engrave, weld, drill, and cut various materials by photochemical, photo-heating, photo-ablating, and photo-mechanical effects, laser processing has become an important and indispensable technique in the modern processing field.

Imaging or processing in turbid media is very difficult compared to in transparent media (e.g. clear atmosphere). The main reasons for these difficulties are two, namely, that the turbid medium strongly absorbs the photoacoustic wave field for imaging or processing, and that these absorbs attenuate the energy of the photoacoustic wave field, and that the turbid medium strongly scatters the photoacoustic wave field for imaging or processing, and that these scatters produce a large amount of photoacoustic noise to drown out the photoacoustic signal produced by imaging, or that the accuracy required for processing is reduced.

The effects of absorption and scattering by turbid media are described by way of example with respect to biomedical imaging. Optical imaging is of particular advantage in the field of medical diagnostics because it has a higher resolution than acoustic imaging and X-ray imaging and also has a higher safety than X-ray imaging (X-rays have a certain radiation damage to human tissue). The absorption and scattering of light waves by human tissue depends on the type of specific tissue. Is related to the composition, refractive index, frequency characteristics, etc. of the tissue. Scattering is also related to the degree of non-uniformity in the refractive index of human tissue, the shape, scale and density of scattering particles, etc., and is therefore a complex function of various parameters (see, e.m.c. hillman, "Experimental and theoretical investigations of Near infrared tomographic imaging methods and clinical applications," Thesis for the degree of Doctor, department of Medical Physics and Bioengineering, university College London, feb.2002; j.l.samdell and t.c. zhu, "a view of in-vivo optical properties of human tissues and its impact on PDT," J Biophotonics,4 (11-12), 2011, pp.773-787). These absorptions and scattering will cause the intensity of the beam to decay rapidly. For example, when near infrared light is applied λ=800 nm) in human blood, the light power decays to 7x 10 of the original incident value after passing a propagation distance of 25 mm under the dual actions of absorption and scattering ^-19 That is, if the original incident optical power is 1 milliwatt, after 25 millimeters of travel in the blood, the optical power decays to only less than 3 photons per second. Of course increasing the incident light power can increase the number of photons last remaining, thereby increasing the signal light intensity required for the last imaging. However, if this is done, the high incident light power will generate more optical noise, and the high incident light power will damage the skin and subcutaneous tissue of the human body.

Underwater imaging, such as in the ocean or in generally more turbid inland rivers, also faces similar problems, although the absorption and scattering rates of general water bodies are much smaller than those of human biological tissues. Because the attenuation of the high-frequency sound wave in the water body is too large, the underwater imaging distance of the high-frequency sound wave is too short, and the imaging resolution of the medium-low frequency sound wave is too low, so that the underwater imaging of the sound wave has no practical significance, and the underwater imaging mainly depends on optical imaging. The absorption and scattering of light waves by turbid water bodies also depends on specific water body conditions, including the composition and proportion of dissolved and suspended matter in the water body, the frequency characteristics of the refractive index of the water body, the shape, scale and density of scattering particles in the water body, etc., and even on the depth of the water, since the composition or density of suspended matter in the water body generally varies with the depth of the water. Therefore, the absorption and scattering of light waves by turbid water bodies are also complex functions of various parameters (see literature: C.D. Mobley, "The Optical Properties of Water," Chapter 43,Handbook of Optics,Vol.I,2ded,Edited by M.Bass,et al,McGRAW-Hill, new York,1995, pp.43.1-43.56; W.Hou, "Active Underwater Imaging," Chapter 4,Ocean Sensing and Monitoring:Optics and Other Methods,SPIE Press Book,2013,p.87-93). Absorption and scattering by the body of water also causes rapid attenuation of the intensity of the propagating beam, for example, even in clear seawater with little turbidity, the furthest propagation distance of light (λ=550 nm) with good transmission is typically only 25 meters. The current furthest underwater optical imaging distance is 4.6 meters to 7.6 meters, and in order to reduce the influence of back-scattered light, the illumination source is preferably placed near the object to be photographed separately from the camera at the time of photographing.

An important problem in laser machining is the depth of the machining. This is because laser machining has a particular advantage in that it can machine an interior region of a material without damaging the surface of the material. For example, the inner region thereof may be selectively etched by focusing a laser beam to the inside of the bulk transparent glass. In addition, in some laser surgery, the laser beam may penetrate the skin of the human body to heat some of the subcutaneous tissues that can efficiently absorb light and heat, such as those containing chromophores, to treat certain diseases by inducing necrosis of the tissues, which is also a treatment technique now called photodynamic therapy.

However, this not only attenuates the light energy of the processing beam due to light absorption and light scattering by the material to be processed, but also causes damage to the internal material in the path of light propagation, and also generates scattered light to reduce the processing accuracy. In addition, in applications where precise positional calibration of the object being processed is required, scattered light can also overwhelm the optical signal of the imaging view. These difficulties have made current laser processing available only for transparent materials such as glass, or for applications requiring only very shallow penetration depths, such as photodynamic therapy of the body's shallow tissues.

Another important application similar to laser processing is the delivery of optical wave energy in turbid media, such as underwater wireless optical communication, atmospheric wireless optical communication, etc., and the problem to be solved is also to increase the efficiency of the delivery of the beam energy in turbid media. Since optical wireless communication has advantages of freedom of movement, high data transmission rate, short time delay, good confidentiality (narrow transmission channel), etc., underwater wireless optical communication has been studied and improved for a long time, devices including a NepTune laser communication system, a single photon detection system, etc. have been fabricated [ see literature: farr et al, "Demonstration of wireless data harvesting from a subsea node using a" ship of opportunity "", oceans 2013,1-5, 2013; and Ya-Ping Li et al, "Experimental realization of underwater optical communication based on single photon detector", proc.SPIE 11763,Seventh Symposium on Novel Photoelectronic Detection Technology and Applications,2021]. However, since water, particularly natural water such as sea water and river water, has strong light absorption and scattering, the power of a light beam is exponentially attenuated in a propagation path, resulting in large energy loss and short transmission distance.

Imaging studies of objects in turbid media have expended a lot of manpower and resources and efforts are continually underway. Related studies have also achieved numerous achievements that have made a great contribution to human society. Taking medical imaging as an example, X-ray fluoroscopy, X-ray computed tomography, ultrasonic inspection and the like have become the most basic and widely-relied tool for modern medical diagnosis due to their non-invasive, high visual and high reliability. In fact, these medical imaging devices have become the basic equipment for supporting modern medical systems, playing an indispensable and vast role for hundreds of millions of people's life health every day.

The study of biomedical imaging, although initiated late, has also emerged a wide variety of techniques and devices, of which there are no depletion of exotic capabilities such as optical coherence tomography devices [ see: J.F.D.Boer, R.Leitgeb and M.Wojtkowski, "Twity-five years of optical coherence tomography: the paradigm shift in sensitivity and speed provided by Fourier domain OCT, "Biomedical Optics Express, vol.8, no.7, 2017, pp.3248-3280]. The difficulty of optical imaging in highly turbid media such as biological tissue in humans is great and the depth of imaging of the human body is still very limited to devices including optical coherence tomography. The maximum optically imaging depth of the human body is only 3 mm, which is currently accepted. However, even at an imaging depth of several millimeters, these optical three-dimensional imaging devices have been widely used in diagnosis and treatment of ophthalmology, dentistry, dermatology, and the like, and have exerted an irreplaceable role.

An incomparable advantage of optical imaging over acoustic imaging and X-ray imaging is that it has a much higher imaging resolution. For example, in the medical diagnosis field, acoustic imaging, X-ray imaging and the like can only see micro blood vessels in millimeter or sublux meter at most, and optical imaging can see cells in micrometer scale, so that optical imaging can implement such fine observation as pathological analysis on diseased tissues of a patient in real time and in real time without painful and time-consuming processes of pathological analysis by firstly making living slices and then making microscopic. Note that the optical imaging resolution: optical coherence tomography resolution of 0.1 μm-1 μm: 1 μm-10 μm, high frequency acoustic imaging resolution: x-ray computed tomography or magnetic resonance imaging resolution of 100 μm: 100 μm to 1000 μm. Optical imaging also has much higher security than X-ray imaging [ see: barton, biomedical Optics, department of Biomedical Engineering, the University of Arizona, tucson, AZ,2015]. Therefore, if the imaging depth of the optical imaging in the human body can exceed 20 cm, since the maximum thickness of the human body is in the chest and generally not more than 40 cm, the imaging depth of 20 cm can realize the optical imaging of the full human body depth when the relative bi-directional imaging is performed from the upper and lower sides of the human body. Once this great goal is achieved, the optical imaging device can replace the now widely used diagnostic devices such as acoustic imaging and X-ray imaging, becoming a supportive technical set of new generation modern medical systems with both high imaging resolution and no radiation hazard.

In the field of underwater imaging, imaging distances of less than 8 meters also pose a great limitation to human underwater observation and activity. Even the most clear seawater severely hinders human vision, making the natural water the largest natural barrier that can mask its internals, including objects that are free to move within it. As human society advances, the underwater activities of humans become increasingly diverse and widespread, and as such, it is eagerly desired to extend underwater vision so as to (at least to some extent) break through this particular natural barrier and greatly increase the freedom of the underwater activities of humans.

Laser processing can handle almost any material, including very hard or very brittle materials, etc. The laser operation can have the advantages of no incision, less bleeding, high operation precision (about 1 micron), high processing speed (shorter than 1 second), no binding and the like. However, current laser surgery is too shallow in depth, such as current photodynamic therapy can only treat tissue within 3 mm of the skin [ see: T.J. Dougherty, et al, "Photodynamic therapy (Review)," Journal of the National Cancer Institute, vol.90, no.12, 1998, pp.889-905], severely limits the most particularly valuable important applications of laser surgery. In addition, although the maximum distance of the current underwater wireless optical communication has been increased to about 300 meters, the requirement of the underwater activities of the human beings is still far lower. Therefore, the depth and the distance of laser processing, light energy transmission including laser incision-free operation, underwater wireless optical communication and the like are also of great significance to the technical development of human society.

Disclosure of Invention

The object of the present invention is to devise a new device for optical imaging and processing in turbid media, which device allows to greatly increase the optical imaging and processing distance in turbid media. This novel optical imaging and processing device has been designed with repeated ideas and multiple difficulties in overcoming the criticality. Such a device effectively reduces both light absorption and light scattering of the turbid medium in the path of the light beam in such a way that the object is illuminated and processed by forming an inner thin light layer in the turbid medium, thereby greatly increasing the imaging and processing distance in the turbid medium. The feasibility of the invention has been demonstrated by related theoretical analysis, and mathematical modeling of the physical model of the device has shown that the device of the invention has excellent performance. Such optical imaging and processing devices achieve imaging signal intensities and optical energy delivery efficiencies far greater than those achieved by prior art devices, for example, which can enhance the intensity of the imaging optical signal by more than 600dB, and have the surprising potential to reach 2000 dB. In human tissue (e.g. human blood) the optical imaging and processing depth of such devices can exceed 5cm, especially imaging depths of up to 20cm may be desirable. In clear sea water, the optical imaging distance of such a device can exceed 100 meters and possibly reach surprisingly 2000 meters, and the distance of its underwater optical wireless communication can also approach 2000 meters. In addition, the basic working principle of the device is not only suitable for light wave imaging and processing, but also can be popularized and applied to imaging and processing of sound waves and even X-ray waves.

The invention of said novel internal optical layer imaging and processing device originates from the following inventive idea: in the path of the illumination and processing beam illuminating and processing the target and the signal beam returning from the target, the turbid medium absorbs and scatters light both from the illumination and processing beam and from the signal beam, which causes both attenuation of the intensity of the illumination and processing beam and from the signal beam and optical noise that floods the signal beam. If one can find a special method: the method can make the illumination or processing light beam disappear in the illumination or processing path, so that the turbid medium cannot absorb and scatter light in the incident path; then the illumination or processing light beam appears at the target inside the turbid medium, namely a thin illumination light layer is formed at the target inside the turbid medium to illuminate or process the target; then, the signal beam reflected by the target is vanished in the return path, so that the signal beam passes through the return path without absorption and scattering; finally, the signal beam is reproduced at a certain position of the return path and received by the imaging receiver; this particular approach thus completely eliminates light absorption and light scattering by the turbid medium for the illumination and processing beams and the signal beam, thereby achieving the desired light imaging and processing without light absorption and light scattering. While such an exclusive idea is not possible in reality, such a singular assumption can be made to be highly approximate by taking advantage of the magic effects of multi-beam interference.

The device is designed as follows: first, a dispersion generating device is used to broaden the pulse width of a short or ultra-short optical pulse, that is, to reduce the peak intensity of the optical pulse by using dispersion to cause coherent destructive interference of various optical wave frequency components contained in the short or ultra-short optical pulse, which can be considered as gradually changing a short or ultra-short optical pulse into a group of light beams having low or very low synthetic light intensity, and then allowing the group of light beams to enter a turbid medium. Note that the light absorption and light scattering of the turbid medium is proportional to the combined light intensity of the set of light beams propagating therein, and thus the light absorption and light scattering caused by the turbid medium becomes small or very small due to its low or very low combined light intensity when the set of light beams propagates in the turbid medium. The turbid medium also has chromatic dispersion, which also affects the interference of the multiple beams within the set of incident beams with each other. If at a certain moment the dispersion value generated by the dispersion generating means is exactly equal to the dispersion value accumulated by the turbid medium in the propagation path and both have diametrically opposite signs, the width of the light pulses stretched by the dispersion generating means will be fully compressed again by the turbid medium, thus reproducing a short or ultra-short light pulse at a certain position within the turbid medium due to coherent constructive interference of the multiple light beams, which short or ultra-short light pulse will form an inner thin light layer for illuminating or processing the object.

The inner light layer illumination device for optical imaging and processing in turbid media comprises: negative dispersion generating means for widening (full width of half maximum of) the short or ultra-short optical pulses; an optical element for making the stretched light pulse incident on the positive dispersion turbid medium; an optical element for returning the short or ultrashort signal light pulses reflected from the target in the turbid medium back along the original incident path or along a different path; negative dispersion generating means for compressing the short or ultra-short signal light pulses re-stretched by the positive dispersion turbid medium; an imaging receiver positioned at an image receiving location for receiving the recompressed short or ultrashort signal light pulses.

The device for the illumination of an inner light layer for optical imaging and processing in a turbid medium and the method on which it is based will be described in more detail below.

Emitting a beam set comprising N beams of angular frequency omega by a laser _j (j=0, 1,2,) N. The N beams have the same or substantially the same amplitude and the same or substantially the same polarization state. The angular frequency spacing Deltaomega of any two adjacent frequency beams of the N beams can be Identical or not, it is assumed here that their angular frequency intervals Δω are identical in order to simplify the analysis and calculation below. Furthermore, at a certain time t, the initial phases phi of the N beams _j (j=0, 1,2,) N is zero. To self-mode-locking the output beams of the lasers and to satisfy the above conditions when their polarization direction is polarized by the polarizer [ see document: smith, "Mode-Locking of lasers," Proc.IEEE,58 (9), 1342-1355, 1970]。

The optical fields of the N light beams overlap each other both in the negative dispersion generating means and in the propagation path within the turbid medium, thereby generating multi-beam interference. Negative dispersion generating devices have negative dispersion, while turbid media have positive dispersion (in general, the dispersion of most optical media in nature is positive dispersion). Thus, since the beams have different frequencies and different phase velocities, coherent destructive interference of the multiple beams causes the composite amplitudes of the N beams to become very small in most of the intervals in the propagation path, and thus the composite light intensities of the N beams are greatly attenuated in the propagation path. In general, the number of N should be large so that the resulting light intensity can decay to the desired small value. The number of N may be from 3 to greater than 10 ¹² (see further description below).

In addition, when the N light beams propagate in the dispersive medium, the phase difference between any two of the N light beams adjacent in frequency also gradually changes due to dispersion. Since the angular frequency spacing Δω between any two frequency-adjacent beams is the same, and the initial phase ω of the N beams has been assumed _j (j=0, 1,2,.. N) is zero at a previous moment, so that the phase difference between any two frequency adjacent beams is first gradually changed from zero to a non-zero negative value in the negative dispersion generating means and then gradually changed from a non-zero negative value to zero again in the positive dispersion medium. When at a certain position in the turbid medium the phase difference between any two frequency adjacent light beams becomes zero or approximately zero at the same time, coherent constructive interference of the N light beams is induced, i.e. the amplitudes of the N light beams add coherently, resulting in a maximum of the combined light intensity.If the number of N is large enough, the maximum value of the combined light intensity can become extremely large and the duration of the maximum value of the combined light intensity can be shortened to be very short, thus forming a very thin layer of internal intense illumination or process light in the medium.

The inner light layer illumination device for optical imaging and processing in turbid media and the method based on the same have the following excellent properties:

(1) In most of the propagation path, the combined intensity of the illumination or processing beam is not only much smaller than the total intensity of the original set of incident beams, but even smaller than the intensity of any one of the individual beams in the original set of incident beams. Since the light absorption and light scattering of the turbid medium are both proportional to the intensity of the illumination or processing beam, the light absorption and light scattering of the turbid medium are greatly reduced, which not only greatly reduces the transmission energy attenuation of the illumination or processing beam, thereby illuminating or processing the target with more light energy, but also greatly reduces the light noise caused by light scattering, thereby effectively improving the signal-to-noise ratio of imaging or the useful processing energy-to-useless noise energy ratio.

(2) When the object to be imaged or processed is placed at a position where the maximum of the combined light intensity occurs, the intensity of the signal beam generated by reflection is greatly increased, because the combined light intensity is generally much greater than the intensity of any one of the individual beams in the incident beam group, but also significantly greater than the total intensity of the original incident beam group (because of the coherent constructive effect of the multiple beams, see further below for the description of the calculation results), and because the intensity of the signal light reflected by the object is proportional to the intensity of the illumination beam.

(3) The signal light pulses reflected back by the target, which typically still contain N beam frequency components of the original incident illumination beam set, have their amplitudes significantly and even non-uniformly attenuated during reflection, and their polarization states slightly and even non-uniformly changed during reflection, but the phase relationship between the beams and the rate of change of the phase of each beam remain unchanged. Thus if the path of the signal beam group returning from the target is exactly the path of the original illumination beam group at the time of incidence (i.e. let the signal beam group return in reverse along the original illumination path), the principle of beam reversibility [ see: goodman, "General principles of Geometric Optics," Handbook of Optics, vol.I,2ed,Edited by M.Bass,and et al,McGRAW-Hill, new York,1995, p.1.10], in the case where the inconsistency of amplitude attenuation and the inconsistency of polarization state change are not severe (this is the actual case in most cases), the signal beam group also undergoes multi-beam interference in the return path, so that the optical amplitude of the signal beam group undergoes coherent destructive interference in most of the return path, and coherent constructive interference occurs only at a certain specific position point. This will reduce the absorption and scattering of light by the turbid medium in the return path of the signal beam set, so that not only the light energy of the return signal beam set is preserved, but also the scattered light noise of the return signal beam set can be reduced. In addition, the signal beam groups are also significantly enhanced again by the resultant light intensity at their coherent constructive location points. Typically the imaging receiver is placed at the point of the combined intensity maxima of this signal beam set. In some cases, the return path of the signal beam set from the target may be different from the path of incidence of the illumination beam set, but so long as the conditions of the above-described multi-beam interference can be met, coherent cancellation of the light amplitudes of the signal beam set may still reduce light absorption and light scattering by the turbid medium in the return path, and coherent constructive addition of the light amplitudes of the signal beam set only at a particular point may result in a resultant light intensity of the signal beam set being maximized and received.

The imaging and processing signals generated by the multi-beam interference inner optical layer illumination imaging and processing device are pulse signals. The same light pulses as the above-described short light pulses comprising N beam components can be used generally repeatedly a plurality of times to illuminate or process a target a plurality of times, which can result in a light pulse train consisting of a plurality of processing light pulses, thereby increasing the total light energy of the light processing; this also results in a signal light pulse train consisting of a plurality of return signal light pulses, thereby increasing the total light energy of the received imaging signal.

The polarized light beam may be plane polarized, elliptically polarized or circularly polarized light beam in the above description of the imaging device of the present invention, because the plane polarized, elliptically polarized and circularly polarized light beams all produce multi-beam interference. The polarization state includes the polarization direction of plane polarized light beams, ellipticity of elliptically polarized light beams, and the like. The N light beams can be plane light beams, cambered surface light beams, cylindrical surface light beams or spherical surface light beams, etc. Since in most applications a planar beam is used, the following physical analysis and mathematical calculations will be based on the case of using a planar beam. For applications using curved, cylindrical or spherical beams, similar steps may be followed for relevant physical analysis and mathematical calculations.

The imaging device of the invention is based on the basic working principle and can also be used for imaging and processing acoustic waves and X-ray waves. Because the light, sound and X-ray waves are all oscillating waves that vary in cosine or sine form, i.e. the amplitude variations of these wave fields are all in cosine or sine form, they all interfere with the multi-acoustic wave field or multi-X-ray wave field and produce similar interference results. In other words, coherent cancellation of the multi-acoustic wave field or multi-X-ray wave field interference may be used to reduce the intensity of the composite wave field in the propagation path, thereby reducing absorption and scattering of the medium, and coherent constructive addition of the multi-acoustic wave field or multi-X-ray wave field interference may be used to increase the intensity of the composite wave field at a particular location to illuminate, process and image the object. The imaging resolution or processing precision of the acoustic wave and the X-ray wave is not high, but the penetration rate of the acoustic wave and the X-ray wave to the turbid medium is much higher than that of the optical wave, so the basic principle of the invention has significance in popularization in the fields of acoustic wave and X-ray wave imaging and processing.

Drawings

The apparatus of the present invention and the method upon which it is based will be further described below in connection with the preferred embodiments for practicing the apparatus of the present invention. It is clear that these preferred embodiments are not the only means that can be devised based on the invention. The described embodiments of the present invention can be changed or modified based on the core principle of the inventive device and the method on which it is based and using the knowledge of the optical design known in the prior art, so that the applicant of the present patent reserves all rights to change and modify the embodiments of the inventive device described below.

The foregoing and advantages of the invention will be apparent from the following description of preferred embodiments of the device and the accompanying drawings (note that, because the size of the various components in the device vary widely, these components are not drawn to scale in order to show the necessary details), in which:

the upper diagram in fig. 1 shows the case where the pulse width of one short optical pulse is stretched by only a positive dispersion optical medium, and the lower diagram in fig. 1 shows the case where the pulse width of one short optical pulse is stretched by a negative dispersion generating means before being compressed by a positive dispersion optical medium. In FIG. 1, the solid line, the broken line and the dot-dash line respectively represent the values corresponding to the angular frequency ω ₀ ，ω _j And omega _N-1 Is arranged in the plane of the wave fronts of the three light beams. When these three planar wavefronts travel in a positive dispersive optical medium, because ω ₀ ＜ω _j ＜ω _N-1 The planar wavefront represented by the solid line travels fastest.

Fig. 2 is a schematic optical construction of a preferred embodiment of a medical imaging or laser incision-free surgical device.

Fig. 3 is a detailed schematic diagram of a portion of the structure of fig. 2.

Fig. 4 shows the variation of the combined light intensity of the multi-beam interference in the optically dispersive medium, where the combined light intensity I of the multi-beam interference varies with the parameter K. Note that only n=1000 is taken in the calculation of the light intensity I. The left and right illustrations in the figures show the shape of the output surface of a dispersion compensating prism of a mirrored negative dispersion device in a medical imaging or laser kerf-free surgical device and in an underwater imaging or underwater wireless optical communication device, respectively (see further details below).

FIG. 5 is a graph showing the variation of the refractive index n of human hemoglobin with angular frequency ω, the variation interval of the angular frequency is 1.44997PHz-3.92699PHz (1PHz=10 ¹⁵ Hz), the wavelength interval corresponding to the angular frequency variation is 1300nm-480nm [ data source: bashkatov et al, "Measurement of tissue optical properties in the context of tissue optical clearing," Journal of Biomedical Optics, vol.23 (9)，2018，pp.091416.1-091416.31]. Note that the change curve due to too few n of the measurement data is not smooth.

Fig. 6 is a graph showing the variation of refractive index n of human body fat with angular frequency ω, and the wavelength variation range corresponding to the angular frequency variation is 1300nm to 480nm [ data source: A.N. Bashkatov et al, "Measurement of tissue optical properties in the context of tissue optical clearing," Journal of Biomedical Optics, vol.23 (9), 2018, pp.091416.1-091416.31].

Fig. 7 is a schematic view of the optical structure of a preferred embodiment of an underwater imaging or underwater wireless optical communication device.

Fig. 8 is a graph showing the variation of the refractive index n of purified water with angular frequency ω. Note that the wavelength range of the visible light frequency region is: 400nm-700nm, the corresponding angular frequency ω has a range of 4.71PHz-2.69PHz [ data sources: "CRC Handbook of Chemistry and Physics," 95th edition,edited by W.M.Haynes,D.R.Lide,T.J.Bruno,CRC Press,New York,2014-2015, p.10.245].

Detailed Description

Using one laser to emit N beams simultaneously, the N beams having different angular frequencies omega _j And the same polarization direction, the angular frequency spacing Deltaω of the beams is the same, i.e

ω _j ＝ω ₀ +jΔω，j＝0，1，2，3，…，N-1， (1)

Wherein omega ₀ Are the N different angular frequencies omega _j Is the lowest angular frequency of (a). In general, these conditions can be satisfied when a polarizer is used to polarize the polarization direction of an output beam from a mode-locked laser [ see: smith, "Mode-Locking of lasers," Proc.IEEE,58 (9), 1342-1355, 1970]。

Assuming that all of the N beams propagate along the x-direction, when the wave fields of the beams are superimposed on each other, their composite wave field can be expressed as

In which A _j (x) Is the optical field amplitude of the jth beam. Since the light absorption and scattering change the light field amplitude with the transmission distance x, A _j (x) Is a function of stroke x. t is the time period in which,is the primary phase of the optical field of the jth beam at x=0 and t=0, C is the speed of light in vacuum, c.c. represents the complex conjugate of the first term in the expression.

Amplitude A can be modified by dispersion compensating the gain of the laser cavity using dyes _j (x) Shape of the distribution with the angular frequency of the light field [ see: R.L.Woodward, "Dispersion engineering of mode-locked fiber lasers," J.Opt.20, 033002, 2018 ]. When A is _j (x) When the distribution shape of (a) becomes rectangular, A _j (x) The value no longer varies with j. Then when the N light beams enter the optical medium, the propagation speed of the jth light beam is changed from C to V _j ＝C/n _j Wherein n is _j Is that the medium corresponds to angular frequency omega _j The refractive index of (2) becomes

Definition of DeltaPhi _j To correspond to an angular frequency omega _j And omega _j-1 Phase difference between two light beams of (2)

When delta phi _j When =0, coherent constructive multi-beam interference occurs for the N light beams, and such multi-beam constructive interference generates a maximum of the combined light intensities of the N light beams, thereby generating a short light pulse with very high light intensity.

This is the case when a short optical pulse emitted by a mode-locked laser enters the input surface of a dispersive optical medium. Since the short light pulse comprises a plurality of light wave components of different frequencies, i.e. N light of different frequenciesA beam, thus corresponding to an angular frequency omega, when an optical pulse enters an input surface of an optical medium _j And omega _j-1 The phase difference between the two beams of (j=0, 1,2,3, …, N-1) is zero. Here, the dispersion effect of the air gap between the mode-locked laser and the optical medium is neglected because the dispersion coefficient of air is very small. Assuming that the input surface of the optical medium is planar and perpendicular to the x-direction and that the input surface of the optical medium is at a position of x=0, if the short light pulse is an input plane into the optical medium at t=0, the initial phases of the N light beams are therefore Zero at x=0 and t=0.

In most cases, the optical medium in nature is a positive dispersion medium, which includes biological tissue and natural seawater in the human body. When an optical pulse enters a positive dispersive medium, beam components of different frequencies contained in the optical pulse propagate at different speeds, and the higher the frequency of one beam, the slower its propagation speed, so the pulse width of the short optical pulse is continuously widened. This process can also be seen as a process in which a short light pulse with a very high light intensity is gradually changed into a beam group with a decreasing combined light intensity, as shown in the upper diagram of fig. 1.

In the upper diagram of fig. 1, N beams come from one mode-locked laser 2, and solid lines, broken lines and dash-dot lines in the diagram respectively indicate that the angular frequency corresponds to ω ₀ ，ω _j And omega _N-1 Is arranged in the plane of the wave fronts of the three light beams. If these three wavefronts travel in the positive dispersive optical medium 4, because ω ₀ ＜ω _j ＜ω _N-1 The planar wavefront represented by the solid line propagates fastest.

While when the three planar wavefronts propagate in the negative dispersion generating means 8, their propagation speeds are opposite, i.e. the higher the frequency of one beam, the faster it propagates. Assuming that a short optical pulse is coming from a mode-locked laser 6, if the optical pulse enters the negative dispersion generating means 8 at t=0, then at position x=0, the three plane wavefronts are mutually Overlapping. Due to omega ₀ ＜ω _j ＜ω _N-1 In the negative dispersion generating device 8, the plane wavefront represented by the broken line propagates fastest.

Thus, in the dispersion generating device 8, although ω _j ＞ω _j-1 ，Δφ _j The value will change from zero to a negative value as the value of x changes. If corresponding to an angular frequency omega _j The propagation distance of the beam in the negative dispersion generating means 8 is x _j Then there is

Multiple beam coherent destructive interference will occur in the negative dispersion generating device 8 with delta phi _j The value increases continuously negative starting from zero.

Note that in the same optical medium, the shorter the original duration of a pulse of light, the faster the pulse is widened and the faster the peak intensity of the pulse of light also decreases. This is because shorter light pulses have a wider frequency range and thus also contain more light frequency components. If the peak intensity of the light pulse is reduced from 100% of the maximum value to a significantly smaller percentage, the length of the fall time of, for example, 0.1% of the maximum value is defined as the pulse widening time T _ib Then the pulse is widened by a distance D _ib ＝V _a T _ib (wherein V _a The average speed of the N light beams) because the peak intensity of the light pulse has been significantly reduced, the light absorption and light scattering of the turbid medium by it becomes small. In practical applications, the desired pulse width time T _ib Or pulse widening distance D _ib Depending on the light absorption and the size of the light scattering coefficient of the medium concerned. For turbid media with greater light absorption or/and light scattering, the desired T _ib Or D _ib The value should be smaller. For example for medical imaging, D _ib The value should be in the order of millimeters. While for underwater imaging, D _ib The value may then be of the order of meters. Also, it can be defined that the peak intensity of the light pulse increases from some small percentage value of its maximum, e.g. 0.1%, to its maximumThe time length of 100% of the value is the pulse compression time T _ls . Since the compression process of the light pulse is completely the reverse of the widening process of the light pulse, T is the same optical medium _ls The value should be equal to T _ib The values are equal.

The widened light pulse leaves the negative dispersion generating means 8 and enters the positively dispersed optical medium 10. In the optical medium 10, the phase difference between any pair of two adjacent light beams of the light pulse is gradually reduced. The optical path difference between any two of the three planar wavefronts shown in the lower graph of fig. 1 will also gradually decrease, so in other words, after entering the optical medium 10, the widening of the pulse width of the light pulse will stop and the compression of the pulse width of the light pulse will start.

Assuming that in a turbid medium it corresponds to an angular frequency omega _j And omega _j-1 The phase difference between the two beams is delta phi' _j Corresponding to an angular frequency omega _j The propagation distance of the light beam in the turbid medium is x' _j Then, the first and second data are obtained,

wherein n' _j And n' _j-1 Is that the turbid medium corresponds to an angular frequency omega _j And omega _j-1 Is a refractive index of (c).

If the medium used in the negative dispersion generating device is the same as the material of the turbid medium or the medium used in the negative dispersion generating device has the same or approximately the same dispersion characteristics as the material of the turbid medium, n' _j ＝n _j And n' _j-1 ＝n _j-1 . Under this condition, if x 'is taken' _j ＝x _j We obtained Δφ' _j ＝-Δφ _j . Thus, since there is Δφ 'for each pair of two frequency adjacent beams' _j -Δφ _j =0, the previously widened light pulse will be fully compressed, thus forming a very thin inner light layer in the turbid medium 10. That is, by implementing the aboveMirror negative dispersion compensation we can form a thin illumination or processing light layer inside the turbid medium as desired.

Fig. 2 shows an optical block diagram of a preferred embodiment of the medical imaging or laser incision-free surgical device as designed.

After passing through the first beam splitter 14, 10% of the light energy from the N parallel and polarized beams from the mode-locked laser 12 enters the second beam splitter 16. The second beam splitter 16 has a light intensity transmittance of 50%. Then, 5% of the total incident light energy is incident on two lenses 18 and 20, which demagnify the cross-section of the passing N beams while slightly focusing the N beams. The beam splitter 14 having a low transmittance is used in order to make the optical energy reflection loss small when the signal beam group is reflected by the beam splitter 14 later. Then, the N light beams whose cross-sectional areas are reduced and slightly condensed are normally incident on one right angle prism 22.

The upper apex angle of the right angle prism 22 is β, so that the N light beams having a small beam diameter and slightly converging when they are incident on the rear output surface of the prism 22, the angles of incidence of the center lines of these light beams to the rear output surface of the prism 22 are also β (see fig. 3). When the N beams are refracted by the prism 22, the N beams correspond to an angular frequency omega _j Angle of refraction theta of the centerline of the beam _j Satisfies [ see literature: s. Goodman, "General principles of Geometric Optics," in Handbook of Optics, mcGRAW-Hill,1995, vol.I ]

n _j sinβ＝sinθ _j (7)

Wherein n is _j Is that the prism 22 corresponds to an angular frequency omega _j The refractive index of air is set to be approximately 1. After being refracted by the prism 22, all the light beams enter a thin lens 24.f (f) ₃ Is the focal length of lens 24. To simplify the analysis and calculation, let the corresponding angular frequency be ω _N-1 The center line of the beam of (c) is parallel to the optical axis of the lens 24 and passes through the mirror center of the lens 24. The optical axis of lens 24 is also parallel to and on the X-axis. U is the distance from the intersection of the output surface of the prism 22 and the center line of each outgoing beam to the mirror center of the lens 24. Because U < ")f ₃ According to newton's thin lens formula [ reference: s. Goodman, "General principles of Geometric Optics," in Handbook of Optics, mcGRAW-Hill,1995, vol.I]Corresponding to an angular frequency omega _N-1 And omega _j Included angle θ between the centerlines of the two beams _N-1 -θ _j Is amplified by M times to become theta' _j That is to say

θ′ _j ＝M(θ _N-1 -θ _j ). (8)

Wherein m=f ₃ /(U-f ₃ ). When M is negative, it is imaged as a virtual image.

The N light beams whose refraction angles are enlarged enter the thin lens 26 in the x direction. The optical axes of the lenses 26 are also parallel and lie on the X-axis. f (f) ₄ Is the focal length of lens 26. The N parallel and slightly converging beamlets then all normal incidence to the input face of prism 28, the input face of prism 28 being planar and perpendicular to the X-axis.

In the prism 28, corresponding to an angular frequency ω _j Is H _j And the distance from the input plane to the output surface along the center line is D _j . From FIGS. 2 and 3 we have

When the beam has traveled a distance D in the prism 28 from the input plane, it corresponds to an angular frequency ω _j And omega _j-1 Will produce an optical path difference deltap _j

ΔP _j ＝DΔn _j ，j＝0，1，2，3，…，N-1， (10)

In the middle of

Δn _j ＝n _j-1 -n _j . (11)

If it is desired to compensate for the positive dispersion generated in the turbid medium, it is desired that any two of the N light beams that are frequency-adjacent should generate an optical path difference in the turbid medium that must be equal to the absolute value of the optical path difference generated in the prism 28 but of opposite sign, i.e. that mirror negative dispersion compensation must be generated in the prism 28.

In order to create mirror negative dispersion compensation in the prism 28, the prism 28 must be made of the same optical medium as the turbid medium or the prism 28 must have the same or approximately the same optical dispersion characteristics as the turbid medium. The condition is fulfilled that the first and second derivatives of the refractive index of the prism 28 for the angular frequencies of the N light beams are both required to be the same as the first and second derivatives of the refractive index of the turbid medium for the angular frequencies of the N light beams, i.e.:

the second derivative of the refractive index of the medium with respect to the angular frequency of the beam determines the group velocity at which the optical pulse travels in the medium, which is a substantial requirement for dispersion compensation [ see: R.L.Fork, O.E.Martinez, and J.P. Gordon, "Negative dispersion using pairs of prisms," Opt. Lett.9 (5), 150-152, 1984].

In recent years, it has become relatively easy to find a material whose refractive index changes with angular frequency to meet the above requirements. For example, it has not been difficult to find a material having optical properties that approximate human biological tissue due to recent developments in simulated optical materials and techniques of human biological tissue. In the fields of research of spectroscopy, medical imaging and medical treatment, such simulation materials have been widely used [ see literature: W.Pogue, and M.S. Patterson, "Review of tissue simulating phantoms for optical spectroscopy, imaging and dosimetry," J.biomed.Opt.,11 (4), 02.1-02.16, 2006]. The dispersion, absorption and scattering properties of these simulated materials are highly similar to those of human biological tissue. Of course, the material from which the prism 28 is made has less light absorption and light scattering, and thus reduces the loss of light energy. If the material obtained is not rigid, the prism 28 may be first formed into a transparent container and then filled with the material obtained.

Assume that two correspond to angular frequencies ω _j And omega _j-1 The optical path difference generated by the light beam in the turbid medium is delta P _j ′＝D′Δn′ _j Wherein Deltan' _j ＝n′ _j -n′ _j-1 ，n′ _j And n' _j-1 Is that the turbid medium corresponds to an angular frequency omega _j And omega _j-1 Is D' is the refractive index corresponding to an angular frequency of ω _j Is used for the propagation distance of the light beam in the turbid medium. After satisfying the above-mentioned requirements for the light dispersion characteristics of the turbid medium, it is necessary for the negative dispersion generating device to generate the following optical path difference

ΔP _j ＝-ΔP′ _j ＝-D′Δn _j ，j＝0，1，2，3，…，N-1， (14)

In addition, in addition to the prism 28, positive dispersion is also generated by other optical elements used in the negative dispersion generating device. If the total length of the propagation path of each beam in the other optical element is much smaller than D', the positive dispersion effect additionally generated in the other optical element can be ignored. Otherwise, these additionally generated positive dispersions also need to be compensated for.

Referring to FIG. 2, assume that the beam has a propagation path length D in the beam splitters 14, 16 and prism 22, respectively _b1 ，D _b2 And D _P1 . Assume again that lenses 18, 20, 24 and 26 are all thin lenses, i.e. gaussian lenses [ see reference: goodman, "General principles of Geometric Optics," Handbook of Optics McGRAW-Hill,1995, vol.I.]Light rays emanating from the focal point of each lens, whether they pass through the center of the lens or through the edges of the lens, become parallel rays after passing through the lens and form a planar wavefront perpendicular to the optical axis of the lens. In other words, all rays emanating from the focal point of each lens will have the same optical path length after passing through the lens. Thus, when the thickness of the thin lenses 18, 20, 24 and 26 in the direction along their optical axes is D _L1 ，D _L2 ，D _L3 And D _L4 Then D _L1 ，D _L2 ，D _L3 And D _L4 I.e. the path length of the light beam through the lenses. Thus, D' will become D

D″＝D′+D _b1 +D _b2 +D _P1 +D _L1 +D _L2 +D _L3 +D _L4 ， (15)

Note that in formula (15), it has been assumed that the beam splitters 14, 16, the prism 22 and the thin lenses 18, 20, 24 and 26 are all made of the same medium as the turbid medium, or they have the same or approximately the same light dispersing properties as the turbid medium.

Since the prism 28 can only be made of a positive dispersive medium material, when ω _j ＞ω _j-1 When it is required that any pair of the pair corresponds to the angular frequency omega _j And omega _j-1 The only way to create a negative optical path difference of less than zero is to change the propagation distance difference of the two beams of any pair in prism 28.

If corresponding to angular frequency omega _j And omega _j-1 The propagation distance of the two beams in the prism 28 is D respectively _j And D _j-1 D is then _j And D _j-1 The following conditions must be met:

D _j n _j -D _j-1 n _j-1 ≈ΔD _j n _j ＝ΔP _j ＝-ΔP′ _j ＝-D″Δn _j ，j＝0，1，2，3，…，N-1. (16)

in DeltaD _j ＝D _j -D _j-1 。

Rearrangement of (16), we have

Equation (17) gives the angular frequency ω in the prism 28 _j And omega _j-1 A negative optical path difference that needs to be generated for the two beams of (a). Thus, for angular frequencies ω _j The total propagation distance D required for the light beam of (c) in prism 28 _j Is that

D _j ＝ΔD _N-1 +ΔD _N-2 +ΔD _N-3 +…+ΔD _j . (18)

Negative dispersion generating devices for light have emerged for decades, mainly for compressing the width of short light pulses widened by a positive dispersion medium [ see: R.L.Fork, O.E.Martinez, and J.P. Gordon, "Negative dispersion using pairs of prisms," Opt. Lett.9 (5), 150-152, 1984 ]. Thus, the use of a typical negative optical dispersion generating device is the exact opposite of our use here. In addition to this essential difference, the existing negative dispersion generating device, either composed of prisms or gratings, cannot generate the sum (n) required by the equation (17) _j-1 -n _j )/n _j Is proportional to the change in optical path length difference. The reason for this is that the negative dispersion produced by the existing negative dispersion generating device is a function of the refraction angle or diffraction angle of the light beam after it is refracted by the prism or grating in the device. For example, the optical structure of the conventional negative dispersion generating device composed of prisms is substantially the same as that shown in fig. 2, but the rear output surface of the prism corresponding to the prism 28 for generating dispersion compensation is a plane. Thus, if each beam is returned from the output surface of the dispersion compensating prism corresponding to prism 28, D _j The value will follow H _j The value varies linearly. Since according to formula (9), H _j The value is a function of the angle of refraction angle, so that the negative optical path difference produced by such a device becomes a function of the angle of refraction angle, which causes distortion in dispersion compensation. This distortion of dispersion compensation becomes severe when the optical pulse width is very short and thus the spectral frequency range involved is very wide (see further discussion below).

The present invention thus also provides an apparatus that produces mirror negative dispersion. The design of the device uses computer controlled high precision optical processing techniques that have evolved significantly in recent years, as well as micro tetrahedral optical retro-reflector techniques that have improved performance significantly in recent years.

The method is realized by the following steps: the refractive index of the desired turbid medium corresponding to the different optical frequencies in the desired frequency range is first measured. Because of the large number of frequencies to be considered, only a part of discrete frequency-dependent refractive index data is needed to be measured in actual operation, and then a computer is used to fit a curve of refractive index variation with frequency in a required frequency range according to the obtained discrete data. Several dispersion equations are available for fitting refractive index profiles, such as the Cauchy, hartmann, conrady and Kettler-Drude equations et al [ see: W.J. Smith, "Optical Materials and Interference Coatings," in Modern Optical Engineering, mcGRAW-Hill,2000,Chapter 7,p.176].

Then, based on the fitted curve of refractive index with frequency and the required D' value, ΔD is calculated according to equation (17) _j Values. Since D' is a known preset imaging or processing distance in the turbid medium and the optical parameters of the optical element used in the negative dispersion generating device are also known, the value of D "can be obtained according to equation (15). Finally, obtaining the angular frequency omega by a computer according to the formula (18) _j The total propagation distance D required for each beam of light _j 。

From formulas (8) and (9), we have

H _j ＝f ₄ tg[M(θ _N-1 -θ _j )]. (19)

From equation (7), we have θ _j ＝n _j arcsin beta and theta _N-1 ＝n _N-1 arcsin beta, thus

H _j ＝f ₄ tg[M(n _N-1 arcsinβ-n _j arcsinβ)]. (20)

Using H _j The value is taken as the position of the j-th point on the Z-axis and is used with H _j Corresponding D _j The value is taken as the position of the point on the X-axis, and the positions of the N points can then be generated in such a way that a data set containing N data pairs (D _j ，H _j ) Is a data set of the data set of (a). The N point locations (D _j ，H _j ) Smoothly connected together to fit a smooth curve. Finally, the rear output surface of prism 28 is machined to the resulting fitted smooth curve shape using computer-controlled high precision grinding and polishing techniques.

In recent years, computer-controlled high-precision optical lapping and polishing techniques have been remarkably developed, which can manufacture even optical components of very large size. For example, in manufacturing a large optical reflector having a radius of curvature of 36m, the manufacturing variation of the surface is less than ±1/36000, and the surface roughness is less than 1nm [ see: D.W.Kim, H.M.Martin, and J.H.Burgea, "Calibration and optimization of computer-controlled optical surfacing for large optics," Proc.SPIE,8126, 15.1-15.10, 2011].

This means that if the optical path difference to be compensated by the prism 28 is 0.1m (which is already a very large optical path difference compensation value, see further below), the maximum machining deviation of the surface shape is + -2.78 μm. The process deviation of the surface shape is much smaller than the optical path difference compensation deviation caused by the dispersion compensation distortion of the existing negative dispersion generating device (see also the further description below).

When N light beams are incident on the output surface of the prism 28 capable of performing mirror dispersion compensation as described above, since the shape of the output surface is generally non-planar and non-circular, the incident angles of the respective light beams are different. In addition, since each incident beam is again a slightly converging beam, even each beam has a different angle of incidence when different rays therein are incident on the output surface of the prism 28. Thus, if the output surface of prism 28 is simply made as a generally reflective surface, these beams are not able to be returned by general reflection back along their original direction and path of incidence.

This problem can be solved by using micro tetrahedral optical retroreflective prisms. Optical tetrahedron is a well known optical retroreflective element that can be seen as a tetrahedron angle cut from one corner of an optical cube, and is therefore also called corner reflector. The optical tetrahedron has 4 equilateral triangular planes, three of which are mutually perpendicular to each other and constitute the light reflecting surface. When a light beam enters such a corner reflector from the fourth face, the incident light beam will return completely in its direction of incidence after being reflected by the 3 reflection planes, regardless of the direction of incidence of the light beam. In recent years, with miniaturization of optical tetrahedrons, a large number of tiny optical tetrahedron prisms are arrayed on the surface of a rigid or soft sheet to make an optical retroreflective sheet. One widespread use of such optical retroreflective sheeting is as a traffic and security marking in nighttime or dim environments. Some optical retroreflective sheeting can retroreflect an incident light beam over a wide spectral range, and can have a reflectivity of greater than 90% when the incident angle of the light beam is less than 30 °. The average diameter of the microprisms may be less than 45 μm [ see: A.Lundvall, F.Nikolajeff, and t.lindstrom, "High performing micromachined retroreflector," opt.express,11 (20), 2459-2473, 2003; A.Poscik, J.Szkudlarek, and G.Owczarek, "Photometric properties of retroreflective materials in dependence on their structure and angle of illumination," fibers text. East. Eur.3 (135), 58-64, 2019].

When a flexible micro-prismatic optical retroreflective sheeting is applied to the rear output surface of the prism 28 with optical glue, the N slightly converging light beams having different angles of incidence can be reversed back along the original direction and path of incidence. In order for the angles of incidence of the different beams to be less than 30 deg. to achieve high reflectivity, the curvature of any point on the output surface of the prism 28 cannot be too small, which can be achieved by increasing the angular magnification M of the lens 24 or/and the focal length f of the lens 26 ₄ To realize the method. Note that the returned beam will again reverse along its original incident path, so D "in equation (15) should be reduced to 0.5D".

The left-hand inset in fig. 4 is the shape of the output surface of prism 28 of the device for medical imaging of the human body or laser incision-free surgery obtained by calculation. In the calculation, H _j And D _j The values of (2) are calculated according to equations (19) and (18). Other parameters used in the calculation are: beta=29.5°, the wavelength ranges from 400nm to 1400nm, and the corresponding angular frequency ranges from 1.346PHz to 4.736PHz (1 phz=10 ¹⁵ Hz). The corresponding refractive index variation ranges from 1.4356 to 1.41, f ₄ =0.5 m, m=1, and D "=30 cm. The desired imaging or laser surgery distance in the human body is 20cm. FIGS. 5 and 6 are folds of human hemoglobin and human adipose tissue The curves of emissivity versus angular frequency of light, which are used in the calculations described above [ see: A.N.Bashkatov, K.V.BereziN, K.N.Dvoretskiy, M.L.Chernavina, E.A.GeNiNa, V.D.GeNiN, V.I.Kochubey, E.N.Lazareva, A.B.Pravdin, M.E.Shvachkina, P.A.Timoshina, D.K.Tuchina, D.D.Yakovlev, D.A.Yakovlev, I.Y.YaNiNa, O.S.Zhernovaya, and V.V.Tuchin, "Measurement of tissue optical properties in the context of tissue optical clearing," J.biomed.Opt.23 (9), 1416.1-1416.31, 2018]。

As can be seen from the left inset in fig. 4, the prism 28 has a height 692mm and a width from 50mm to 52.54mm. In practice, the maximum optical path compensation in the prism 28 is only 2.54mm, so the width of the prism is made to exceed 50mm in order to give the prism 28 sufficient mechanical strength. In order to obtain a sufficiently short widening time T in the human body _ib An ultra short light pulse of about 2fs is used. The calculated shape of the output surface of the prism 28 is clearly non-planar and non-circular, which illustrates that if the output surface shape of the prism 28 is planar as the prisms used in existing negative dispersion devices, the desired mirrored negative dispersion compensation cannot be obtained.

There is a need to further discuss the dispersion compensation errors present in the designed negative dispersion generating device. For most of the N slightly converging beams, their focal points do not fall exactly on the output surface of the prism 28. In addition, for each beam, only the central ray of the beam can precisely propagate D in the prism 28 _j Thus this all causes errors in optical path compensation. But it can be demonstrated that the induced optical path compensation error is not greater than 25 μm. FIG. 3 shows a frequency corresponding to an angular frequency ω ₀ ，ω _N-1 And omega _c Is a light path diagram of three slightly converging light beams. If the diameter of each slightly converging fine diameter beam on the input surface of the prism 22 is delta, delta may be less than 1mm for the visible light beam without problems. F is the focal plane of the N light beams after being focused by lens 20 (note that the converging action of lens 24 and lens 26 on the N light beams is much smaller than that of lens 20, and thus the converging action of lens 24 and lens 26 onThe focusing effect of the N beams is ignored in this analysis), the F-plane is set perpendicular to the X-axis, and passes through (D on the output surface of the prism 28 _C ，H _C ) And (5) a dot. (D) _C ，H _C ) The point corresponds to an angular frequency omega _c Note ω _c ＝(ω _N-1 -ω ₀ )/2. Thus, the diameter of the N beams on the output surface of the prism 28 should be much smaller than the delta value. For example, if the distance along the X-axis from the mirror center of the lens 20 to the plane F is L _C And if D ₀ -D _N-1 Is L _f When δ=1 mm and L _f ＝L _c At/20, because plane F is at D ₀ And D _N-1 In the middle of the two distances, the diameter of the N beams on the output surface of the prism 28 should be smaller than 25 μm, so that the optical path compensation error caused by the non-uniformity of the beam diameters on the output surface of the prism 28 will be smaller than 25 μm.

The optical path compensation error caused by the microprism optical retroreflective sheeting is also small. As described above, the average diameter of the microprisms is less than 43 μm, so if the average depth of the microprisms is assumed to be half of 43 μm, the optical path error caused by reflection of the thin mirror layer of the microprisms should be less than half of 43 μm, considering that an approximately planar wavefront is generated when rays of the light beam return from each of the retroreflective microprisms.

The N beams returned from prism 28 are recombined in prism 22. After being reflected by beam splitter 16 and mirror 30, these beams become parallel beams and enter an imaging or machining distance adjuster consisting of two triangular elements 32 and 34. The imaging or processing distance adjuster is required because after the size of the prism 28 and the shape of the rear output surface are determined, the imaging or processing distance D 'of the relevant device in the turbid medium is determined by equations (15), (17) and (18), i.e. each fabricated device has a fixed imaging or processing distance D'. Thus, if the desired imaging or machining depth in the human body is D _b The distance D is changed by sliding the position of the two triangular elements 32 and 34 ₃ And D ₄ Then make it again

D′＝D ₃ +D ₄ +D _b . (21)

Desired imaging or machining depth D _b Can be adjusted by adjusting D ₃ +D ₄ Is changed by the value of (a). Note that the two elements 32 and 34 should also be made using the same medium material as the turbid medium, or the medium material from which the two elements 32 and 34 are made should have the same or approximately the same dispersion characteristics as the turbid medium. Because the two triangular elements 32 and 34 have symmetrical shapes, the imaging or machining distance adjuster does not produce unwanted additional dispersion.

Then, the N parallel light beams enter the human body 40 by reflection of the reflecting mirror 36. Without lens 38, the N beams would have a depth D in the human body _b An interior illumination or processing light layer is produced. The thickness of the illuminating or processing light layer is determined by two factors, one being how many beams N are and the width Δv of the total frequency range of the N beams (2pi Δv=Δω). According to the principle of uncertainty, there is nΔvΔτ=1 (where Δτ is ΔΦ) _j Duration width of light pulse at=0) [ see document: W.H. Carters, "Coherence theory," in Handbook of Optics, mcGRAW-Hill,1995, vol.I, p.4.3 ]Thus, δh=v can be used _h The relationship of Δτ determines the thickness δH of the interior illumination or processing light layer, where V _h Is the average speed at which the N beams propagate in the human body. If the total spectral width of the N beams is sufficiently wide, the thickness of the illumination or processing light layer may be very thin, e.g. may be less than 1 μm.

Another factor is the broadening time T, which is determined by the dispersion characteristics of the medium _ib Is a length of (c). The broadening rate of the optical pulse width caused by medium dispersion can be estimated as follows [ see: C. A.Bunge, M.Beckers and B.Lustermann, polymer Optical Fibres, fiber Types, materials, diagnostics, characterization and Applications, elsevier Ltd, woodhead Publishing,2017, pp.47-118]

ΔT′＝L′Δλd _c ， (22)

Where Δλ is the spectral width, d, of the optical pulse calculated in terms of wavelength _c Is the dispersion coefficient of the turbid medium, L' is the propagation distance of the optical pulse in the dispersive mediumFrom, ΔT' is the full time width (FWHM) of the half maximum intensity value of the light pulse. For example, for a general optical communication fiber, d _c Is typically 20 ps/nm-km at a center wavelength of 1550 nm. Thus, if Δλ=1000 nm, for a light pulse of 2fs, Δt '=20fs when L' =1 mm. When the width of the 2fs optical pulse is widened to 20fs, the peak optical intensity of the optical pulse should drop below 10% of its maximum value.

For sea water, typical d _c Values from 60 ps/nm.km to 300 ps/nm.km [ see: "Seawater intrusion and mixing in estuaries," Marine Species Introduced Traits Wiki,2020, marinespecies. Org/introduced/wiki/seawater_interval_and_timing_in_estuaries# exact details_of_the_longitudinal_dispersion_coeffcient]. Since the value of the dispersion coefficient of the human biological tissue is temporarily not found, taking into account that 60% of the human weight is water, the value of the dispersion coefficient of the sea water is temporarily used here to make a preliminary estimation by approximating the mean value of the dispersion coefficient of the human tissue. Approximate calculations show that if an ultrashort optical pulse of the order of femtoseconds is first broadened by a negative dispersion generating means, then the compression time T is due to it _ls Very short, ultra-short light pulses on the order of femtoseconds can be compressed in human tissue sufficiently fast. Therefore, the light energy loss due to light absorption and light scattering is very small in most transmission paths in the human body (because of the compression distance D of the femtosecond light pulse in the human body _ls Very short, on the order of less than a millimeter). Fortunately, it has not been difficult to obtain ultra-fast high power lasers on the order of femtoseconds in the laser market today.

In fig. 2, the N beams are converged by lens 38 to be incident on human body 40, which is to employ a confocal imaging structure to improve the longitudinal resolution of imaging (as will be further explained later).

When the device is used as a laser kerfless surgical device, the output power of the mode-locking laser 12 is increased to the level required to treat the target tissue, which may be irradiated, heated, ablated, vibrated broken, or the formed endo-optical layer is cut or resected, which may be referred to as a laser scalpel. When the working light power is increased to treat the target tissue, the light intensity in the beam propagation path may still be well below the safety threshold for not damaging other tissues in the path due to multi-beam interference (see further description below).

When the device is used as an imaging device, or when the examination of the post-operation result of the target tissue subjected to laser operation is needed, or when the imaging observation is needed to be carried out on the target tissue, after the target tissue is precisely aimed by a laser beam, the target is illuminated by a laser pulse with lower energy, and then the signal light pulse reflected by the target tissue in the human body is reversely returned along the original light beam incidence path. Typically, all N beams of light that make up an incident illumination light pulse will be reflected by the target tissue. The reflection of the signal light pulse occurs at the surface of the target tissue, i.e. at the interface formed by the surface and the adjacent regions of different refractive index (the two regions of different refractive index form an interface between the two regions). The reflectivity of N beams at the same interface is typically not much different. In the return path, since the signal light pulse still contains N light frequency components, i.e. contains a plurality of light beams of different frequencies, the width of the signal light pulse will again be widened by the human biological tissue with positive dispersion, which results in that the combined light intensity of the signal light beam set is again reduced in the return path, so that the light absorption and light scattering of the signal light pulse in the return path is also reduced. The signal beam group then exits the human body into an imaging or processing distance adjuster, and the optical path difference of the frequency adjacent beams in the signal beam group is further increased in the imaging or processing distance adjuster having positive dispersion. After reflection by the beam splitter 16, the signal beam group again enters the prism 28, this time the widened signal light pulse will be compressed via the negative dispersion effect of the prism 28, as in the case of the widened light pulse being compressed in the conventional negative dispersion generating device. Since the return process of the signal light pulse is the exact inverse of the illumination process of the original incident light pulse, a detailed mathematical analysis of the return process of the signal light beam is not performed.

When the signal beam group reaches the beam splitter 14 again and is reflected, and when the signal beam group propagates a distance exactly equal to D ", the expected short signal light pulses will occur by coherent constructive interference of the N signal beams. Because beam splitter 14 has a relatively high reflectivity, a substantial portion of the energy of the signal beam group will be focused by lens 44 onto image plane 48.

The effect of dispersion of the air gap between the optical elements in the device is neglected for reasons already described.

If the lens 38 is not used, the imaging structure of the device is a widely used optical imaging structure which can directly change any point on the object plane into a corresponding point on the imaging plane, and one of the advantages of this imaging structure is that it can be easily combined with super-resolution imaging technology [ see literature: huszka, and M.A.M.Gijs, "Super-resolution optical imaging: a compactison, "Micro and Nano eng.2,7-28, 2019]. For example, placing a bit-phase filter in front of the focusing lens 44 can cause the imaging resolution to exceed the diffraction limit of optical theory, which can cause the lateral imaging resolution at the image plane to be less than the wavelength of the light beam used.

The lens 38 is used to improve the longitudinal imaging resolution of the device. Because of the compression time T when the illumination beam group is compressed stepwise in the human body _ls In this case, the combined light intensity of the beam combination becomes significantly greater. For example, as previously described, if the beam set is transmitted in human tissue, the combined intensity of the beam set corresponding to a light pulse of 2fs has been greater than 10% of its maximum value in this interval when the compression distance is 1 mm. The actual effective light pulse width of the 2fs light pulse when gradually compressed in human tissue will be much larger than its theoretical width in human tissue (about 0.2 μm), which will reduce the longitudinal resolution of the device imaging. Confocal imaging can improve this problem [ see: S.Inoue, and R.Oldenbourg, "microscope," in Handbook of Optics, mcGRAW-Hill,1995, vol.II]. Longitudinal separation of the target during imaging by focusing the illumination beam using lens 38 to scan the target tissue within the body and using a spatial pinhole diaphragm 46 placed in front of the image plane 48 to mask out-of-focus light outside the focal regionThe resolution can be increased to the wavelength order of the light beam used, i.e. around 1 μm, and contributes to an increase in imaging contrast. When the device is used as a laser incision-free surgical device, the imaging elements in the device, including the lens 44, pinhole diaphragm 46, and imaging receiver are all removed.

Fig. 7 is a schematic view of the optical structure of a preferred embodiment of an underwater imaging or underwater wireless optical communication device. It comprises mode-locked laser 60, beam splitters 62 and 64, thin lenses 66, 68, 72, 74, 80, 82, 92, 94 and 100, prisms 70, 76, mirror 78 and spatial pinhole aperture 102. Since the device is constructed in large part as the medical imaging or laser incision-free surgical device described above, we will only describe imaging distance adjusters that differ in structure from the device.

Its imaging or working distance adjuster is composed of two parallel planar mirrors 84 and 86 and two triangular elements 88 and 90. The angle θ between the mirror surfaces of the two plane mirrors 84 and 86 and the Z axis _M 。

Since the distance of underwater imaging or machining is generally long, each of the N parallel beams of light needs to be reflected multiple times in two planar mirrors of the imaging or machining distance adjuster after entering the adjuster. The diameter of each beam is small so that each beam can be reflected multiple times between two mirrors. If the distance to be imaged and processed in the body of water is D _b D' is the intrinsic imaging and processing distance of the device, and the optical path distance change produced by adjusting the imaging or processing distance adjuster is iD _w Let D' =d _b +∏D _w By changing the pi D _w Can change D _b . Where pi is the total number of reflections of each beam back and forth between two planar mirrors, D _w Is the change in optical path length produced by each reflection of each beam in two planar mirrors. D is not depicted in FIG. 6 _w As it is a variable amount with the movement of the two triangular elements 88 and 90. As can be seen from FIG. 7, when the distance between the mirrors is W, the distance that the beam segment moves up the mirror surface per reflection is Wtg (θ _M ) Thus when twoWhen the height of the mirror is H ', the maximum number of times each beam can be reflected between the mirrors is n=h'/Wtg (θ _M ). The maximum distance that each beam propagates between the mirrors isSo a smaller beam incident angle theta _M And a larger mirror height H' may give a greater adjustment of the imaging or processing distance. Since the diameter of each beam may be small, e.g., 5 mm, the thickness of the imaging or machining distance adjuster may be thin, e.g., less than 10 mm, the imaging or machining distance adjuster may have a suitable volume and weight. When H' =1 m, tg (θ _M ) =0.01 sumAt the time of II D _w /Wtg can be 100m. Still further, if desired, the imaging or machining distance adjuster may be a composite distance adjuster composed of a plurality of single imaging or machining distance adjusters, for example, 10 single distance adjusters. In this case, if each distance adjuster is adjusted to have a distance of 100 meters, a height of 1 meter, and a thickness of 10 millimeters, the complex adjuster may have an imaging or processing adjustment distance of 1000 meters, a height of 1 meter, and a thickness of 10 centimeters.

The illumination beam group exits from the imaging or processing distance adjuster, and after the beam is expanded by lenses 92 and 94, if the beam is not converged by lens 94, N beams enter the water 98, and the N beam groups have a distance D in the water due to multi-beam interference _b The resulting intensity of light is extremely high, resulting in a thin illumination layer 96.

When the device is used as an underwater wireless optical communication device, desired underwater wireless communication can be performed by appropriately adjusting the output power of the mode-locked laser 12. The output power of the mode-locked laser 12 can also be increased to perform operations such as irradiation, heating, ablation, etc. on the underwater target. And the lens 100, the spatial pinhole aperture 102 and the image receiving element may be removed.

When the device is used as an underwater imaging device, or when the result inspection is needed to be carried out on an underwater target subjected to laser processing, or when the underwater target is firstly imaged and observed, and the target is precisely aimed by a laser beam and then is subjected to laser processing, the target is firstly illuminated by laser pulses with lower energy, then an illumination light layer formed in a water body illuminates the target object, and signal light pulses generated by reflection from the target are reversely returned along the incidence path of an original illumination light beam group and undergo a process similar to that described in a medical imaging or incision-free laser operation device. Finally, the signal beam group is reflected by the beam splitter 62 and produces the desired short signal light pulses, which are focused by the lens 100 onto the image plane 106.

Likewise, to increase the longitudinal resolution of the resulting image, the N beams may be focused by lens 94 to scan objects in the sea and a spatial pinhole diaphragm 102 placed in front of the image plane 106 to perform confocal imaging.

The right interpolation diagram in fig. 4 shows the shape of the output surface of the negative dispersion generating prism 76 obtained by calculation for use in underwater imaging and underwater wireless optical communication devices. In the calculation, H _j And D _j The values of (2) are also calculated according to equations (19) and (18). The parameters used in the calculation are: beta=29.5°, the wavelength ranges from 546.4nm to 550nm, and the corresponding angular frequency ranges from 3.427PHz to 3.450PHz. The refractive index varies from 1.4356 to 1.4358.f (f) ₄ =10m, m=50, and D "=1000m. The expected imaging distance in seawater is 1000 meters. The calculations indicate that the height required for prism 76 is 76.4mm and the width required is from 1.70mm to 78.1mm. Since light pulses of about 0.5ps duration are used, the corresponding spectral width is only 3.6nm. The output surface shape of the prism 76 is approximately planar. Fig. 8 shows a graph of refractive index of pure water according to calculation as a function of angular frequency of light field [ see literature: W.M.Haynes, D.R.Lide and T.J.Bruno, CRC Handbook of Chemistry and Physics,95 ^th ed，CRC Press，2014-2015，PP.10.244-10.255]。

Next, we give mathematical certainty to the excellent imaging performance of the device of the present inventionAnd (5) calculating and proving the quantity. From equations (1) and (3), and because of the initial phases of the N beamsEqual to zero when x=0 and t=0, we have

In nature, the refractive index of many optical media is approximately linear with the angular frequency of the light wave, particularly water, a property which can be seen in fig. 5, 6 and 8. Note that the measurement of the refractive index of human biological tissue is not easy, so the refractive index-versus-angular frequency curves in fig. 5 and 6 are not smooth due to the lack of measurement data and unavoidable measurement errors. Furthermore, since the total spectral width of the N light beams typically required is not wide, for example for underwater imaging or processing, to simplify the theoretical analysis, we approximately assume that the refractive index of the turbid medium under consideration varies linearly with the angular frequency of the light waves. As previously mentioned, this approximation is acceptable for medical imaging or processing as well as underwater imaging or processing due to the large amount of water present in the human body. Therefore, we can assume that there are

n _j ＝n ₀ +jΔn，j＝0，1，2，3…，N-1， (24)

Wherein n is ₀ Is that the turbid medium corresponds to an angular frequency omega ₀ Is a refractive index of (c). From formula (1), we have

Δn＝kΔω. (25)

Where k is a proportionality constant. Thus, the first and second substrates are bonded together,

because Deltaomega < omega ₀ Even because (N-1) Δω is still smaller than ω in many cases ₀ Therefore, the formula (26) can be approximately written as

The phase term in brackets in formula (23) is replaced with formula (27), and formula (23) becomes

Wherein k=t- (2 xn) ₀ /C)。

The combined light intensity variation of the multi-beam interference of N beams in the dispersive medium can be obtained by utilizing a triangle series summation formula

When kΔω becomes zero, the value of I will become maximum. The results of the numerical calculations made according to equation (29) are shown in fig. 4 and table 1. The curve in fig. 4 shows the variation of the value of the combined light intensity I with the parameter K, and the parameter K is used as a unit of the lateral coordinate to avoid further complicated theoretical derivation, while the plotted curve of the combined light intensity can still show the essential characteristics of multi-beam interference in the optical dispersive medium. The I-curve in fig. 4 can be seen as the variation of the resulting light intensity of the instantaneous multi-beam interference with propagation distance x of the formation of an inner thin light layer in the turbid medium. In the calculation, a=1 is selected. In addition, in order to clearly draw the width of the synthesized light intensity maximum, only n=1000 is taken.

In table 1, N is the number of beams participating in interference. Gamma is the maximum synthetic light intensity gamma I ₀ Is a factor of enhancement of (1). Epsilon is the residual composite light intensity epsilon I between two composite light intensity maxima ₀ Is shown in fig. 4). I ₀ Is the original incident intensity of each of the N beams. In table 1 we see that the gamma value increases rapidly with increasing N. In the beginning of the increase of N, fluctuations in the epsilon value may be caused by insufficient accuracy of the calculation of the computer or software program we use.

Mu is added to _a Sum mu _s Defined as the light absorption and light scattering coefficient of the turbid medium. For light propagation in turbid media, the Beer-Lamber law is followed [ see reference: J.D.J.Ingle, and S.R.Crouch, spectrochemical Analysis, prentice Hall,1988]The light intensity I varies with the propagation distance x

Wherein NI is ₀ Is the total incident light intensity of the N beams. When the propagation distance of the incident beam group is D', it is assumed that it is reflected by the target and becomes a signal beam group. If the reflectivity of the target object is R, the intensity of the returned signal beam set reaches the observer

In equation (31), it has been assumed that all optical elements used in the device have very low light absorption and light scattering, and thus D 'may be replaced by D'. When the device of the invention is used, the combined light intensity of the beam combination is reduced to εNI in the beam propagation path ₀ . Because of mu _a Sum mu _s Respectively, the rate coefficient describing the attenuation of the light intensity per unit distance due to light absorption and light scattering, if the remaining light intensity of the group of light beams is reduced to NI ₀ This corresponds approximately to a reduction of the attenuation rate of the light absorption and light scattering to a factor epsilon of its original rate, so that when the propagation distance of the light beam set in the turbid medium is x, the change in the light intensity of the light beam set due to the light absorption and light scattering can be written as

When the beam group propagates a distance D 'and is reflected by the target object to become a signal beam group, if the reflectivity of the target object is R', when the returned signal beam group reaches the observer, the light intensity of the signal beam group becomes

TABLE 1 optical signal intensity enhancement for multiple beam interference imaging at different depths in the human body

The ratio of I' in formula (33) to I in formula (31) is defined as the signal strength enhancement factor, i.e., the ratio of the signal strength produced by the inventive device to the signal strength produced by a typical conventional imaging device. From the slave

And when the device of the invention is used, R' I is used as a result of ₀ ＝γRI ₀ Then by the finishing formula (34), we have

Where α is a signal strength enhancement factor. Further, by letting I' =ζni ₀ Then from equation (33), we have

Xi is a signal intensity comparison factor, which is the total incident light intensity NI of the signal light intensity I' received by the device and N original illumination light beams ₀ Ratio of the two components. In the following, we will calculate the value of the enhancement factor α of the signal light intensity and the comparison factor ζ of the signal light intensity according to the light absorption and light scattering conditions of the human biological tissue and the clear seawater using the above-given calculation formula.

In the human body, differentHas different refractive indices, different light absorption and light scattering coefficients. Therefore, in actual analysis and calculation, only the average value of these parameters can be used for analysis and calculation. Since a large amount of blood and water are contained in the human body, the refractive index, light absorption coefficient and light scattering coefficient of blood are temporarily used as the average value of various human tissues to calculate the α and ζ values. The light absorption and scattering coefficient of human blood is mu _s ＝0.397mm ^-1 Sum mu _s ＝1mm ^-1 See literature: D.J.Faber, C.G.Aalders, E.G.Mik, B.A.Hooper, M.J.C.V.Gemert, and T.G.V.Leeuwen, "Oxygen saturation-dependent absorption and scattering of blood," Phys.Rev.Lett.93 (2), 8102.1-8102.4, 2020]. Details of the target biological tissue within the human body are not depicted in fig. 2. In fact, various tissues in the human body fill the entire body and are connected to each other. The illumination light of the internal illumination light layer formed in the human body is reflected at the interface between two regions having different refractive indices. Here, we assume that light is reflected from the interface between blood and fat, and because the refractive indices of human blood and fat are 1.3771 and 1.4714, respectively [ see document: A.N.Bashkatov, K.V.BereziN, K.N.Dvoretskiy, M.L.Chernavina, E.A.Genina, V.D.Genin, V.I.Kochubey, E.N.Lazareva, A.B.Pravdin, M.E.Shvachkina, P.A.Timoshina, D.K.Tuchina, D.D.Yakovlev, D.A.Yakovlev, I.Y.Yanina, O.S.Zhernovaya, and V.V.Tuchin, "Measurement of tissue optical properties in the context of tissue optical clearing," J.biomed.Opt.23 (9), 1416.1-1416.31, 2018 ]According to the fresnel formula [ see: J.M. Bennett, "Polarization," in Handbook of Optics, mcGRAW-Hill,1995, vol.I]We have found that the reflectance of the illumination light is r=0.0011. Finally, we derive from equations (35) and (36) the values of the intensity enhancement factor α and the intensity comparison factor ζ of the intensity of the signal reflected from the target tissue at different depths (D' values are 2cm,5cm,10cm,15cm, and 20cm, respectively) in the human body, based on the values of the intensity enhancement factor γ and the intensity attenuation factor ε shown in Table 1, which values are also shown in Table 1.

For underwater imaging, take 0.0196m ^-1 And 0.0212m ^-1 Light absorption coefficient mu as transparent seawater _a And light scattering coefficient mu _s See literature: C.D. Mobley, "The Optical Properties of Water," in Handbook of Optics McGRAW-Hill,1995, vol.I]. Assuming that the refractive indices of seawater and the target object are 1.34 and 1.6, we obtain that the reflectivity of the target object is r= 0.00809. Then, from the values of the signal light intensity enhancement factor γ and the light intensity attenuation factor ε shown in Table 2, from formulas (35) and (36), the values of the optical signal intensity enhancement factor α and the optical signal intensity comparison factor ζ at different distances D '(D' equal to 200m,500m,1000m,1500m, and 2000m, respectively) in sea water can be obtained, which are also shown in Table 2.

TABLE 2 optical signal intensity enhancement for multiple beam interference imaging at different distances in seawater

As can be seen from tables 1 and 2, the maximum value of the synthesized light intensity of the signal light is much higher than that obtained by the ordinary imaging device. For example, for medical imaging or processing, when the imaging or processing depth is D '=5 cm, or for underwater imaging or processing, when the imaging or processing distance D' =500 m, the enhancement factor ζ of the signal light intensity may be greater than 1.1x10 ³ . This means that when N > 10 ³ When the total spectrum width of the N light beams is wide enough, the peak intensity of the signal light pulse can be higher than NI ₀ . Note that NI ₀ Is the total beam intensity when the N beams are incoherent beams. Of course, extremely high peak intensities of the light pulses are often accompanied by extremely narrow light pulse durations, and thus the light energy per light pulse may be very low. But the desired processing or signal light energy can be obtained by applying or receiving a plurality of repeated signal light pulses as long as the signal-to-noise ratio of the optical signal is high. As can be seen from tables 1 and 2, even when the imaging or processing depth D' =20 cm is formed or processed in medical imaging or processing, or when imaging or processing is performed underwater, as The signal light intensity enhancement factor ζ value is still high like or the processing distance D' =2000 m. Thus, there is an excellent possibility to obtain a signal strength enhancement factor approaching 2000dB using the imaging device of the present invention, which corresponds to an imaging or processing depth of 20cm in humans and an imaging or processing distance of 2000m in clear seawater. Considering that there are some approximations in the calculation, the signal strength enhancement factor of the imaging device of the present invention is represented by a more conservative 600dB, which corresponds to an imaging or processing depth of 5cm in the human body and an imaging or processing distance of 1000m in clear seawater. It can also be seen from fig. 4, tables 1 and 2 that when the device is used for processing objects in turbid media, the combined light intensity value in the beam propagation path will be well below the safety threshold despite the internally formed thin light layer, i.e. the combined light intensity maximum of the illumination or processing beam set, being very high. This is because, as shown in the theoretical calculation results given in tables 1 and 2 of fig. 4, the resultant light intensity maxima may be more than 14 orders of magnitude higher than the resultant light intensity value between the two light intensity maxima, such a large intensity difference providing sufficient space for non-target tissues and objects in the path of the laser beam to avoid photodamage. For example, in laser surgery applications, optical power densities less than 10 milliwatts per square centimeter are safe for human tissue, including skin, while power densities greater than 10 watts per square centimeter ablate most human tissue, the latter being only 3 orders of magnitude higher than the former.

It is not surprising that the device of the present invention can have excellent performance, which is the magic effect of multi-beam interference. By means of multi-beam interferometry, one can now obtain several tens of attosesond=10 ^-18 s) and several watts (1 terawatt=10) ¹² watts), which is the fastest and strongest "artificial event" that humans can produce so far, we believe that multibeam interferometry will make more technical contributions to human society in the future.

Finally we present a practical way how to calibrate the imaging and processing distance of the device in a turbid medium so that imaging or processing can be performed accurately. A turbid medium is typically composed of a plurality of components having different refractive indices, and the imaging and processing distance of a device is determined jointly by the refractive indices of a plurality of refractive index-different cells in the path of the light beam, so that it is difficult to obtain an accurate imaging and processing distance in a specific turbid medium. There is a simple and feasible way to overcome this difficulty. Because in the incident set of N illumination beams, as long as the phase difference between any two adjacent beams becomes zero or an integer multiple of 2 pi, the resultant light intensity generated by coherent constructive interference must appear extremely large, i.e. a thin bright illumination layer must appear inside the turbid medium, so as if one had previously searched for the desired radio station by tuning the frequency of the ordinary radio, you do not need to know what the exact frequency of the station is, i.e. what the exact distance of imaging and processing needed is, you only need to adjust the imaging and processing distance adjuster of the device while looking into the turbid medium, and when the object in the turbid medium appears and becomes clear, the exact imaging and processing distance in the turbid medium that is desired is found.

Pulse broadening is a well-known physical phenomenon [ see: amemiya, "Pulse broadening due to higher order dispersion and its transmission limit," J.Lightwave technology.20 (4), 591-597, 2002], is even a ubiquitous technical hurdle in short-pulse long-haul fiber optic communication systems that degrades communication quality. Negative dispersion generating means have therefore been used to compress broadened short optical pulse signals [ see: R.L.Fork, O.E.Martinez, and J.P. Gordon, "Negative dispersion using pairs of prisms," Opt. Lett.9 (5), 150-152, 1984].

Essentially, the imaging method used by the device of the present invention is a reverse physical process of the process of passive broadening and then artificial compression of short optical pulses that occur in short pulse fiber optic communication systems. In the device of the invention, the short light pulses are artificially widened by the negative dispersion generating device and then automatically compressed by the positive dispersion optically turbid medium. In order to ideally shorten the optical pulse, especially for ultra-short optical pulse, we designed a special mirror negative dispersion generating device.

According to the principle of light reversibility [ see: the feasibility of the imaging or processing method on which the device according to the invention is based is unquestionable, only in terms of how to improve the quality of the internally illuminated or processed light layer produced in the turbid medium by the device according to the invention, since two optical processes in opposite directions must take place in both optical paths if they have mirrored optical structures, as is the case in s.goodman, "General principles of Geometric Optics," in Handbook of Optics, mcGRAW-Hill,1995, vol.i. With the continued improvement in the performance of mirrored negative dispersion generating devices, it is believed that the devices of the present invention will be capable of generating high quality illumination or processing light layers inside turbid media.

The novel optical imaging or processing device created by the invention has important application prospect, and the novel optical imaging or processing device is believed to be widely applied in the future.

Claims (7)

1. An apparatus for illuminating an inner optical layer for optical imaging and processing in a turbid medium, the apparatus comprising: lasers capable of emitting short or ultrashort light pulses containing N optical beam components having different frequencies, typically the same frequency spacing, the same or approximately the same polarization state, and an initial phase that is zero at a previous time; mirror image negative dispersion generating device capable of widening the time width of short or ultra-short light pulse; an imaging or processing distance adjuster capable of adjusting an imaging or processing distance in the turbid medium; a receiver capable of imaging a signal light pulse reflected from a target in a turbid medium.

2. The apparatus for internal optical layer illumination for optical imaging and processing in turbid media according to claim 1 characterized in that said internal optical layer illumination or processing is performed by first using mirrored negative dispersion generating means to widen the temporal width of the short or ultra-short optical pulses; then causing the widened light pulse to enter the turbid medium after passing through the imaging or processing distance adjuster, and compressing the time width of the widened light pulse by using positive dispersion of the turbid medium and the imaging or processing distance adjuster in the propagation path of the light pulse, and reproducing the short or ultra-short light pulse at the target position in the turbid medium to illuminate or process the target; the short or ultra-short signal light pulses reflected by the target are returned reversely along the incident path of the original illumination light pulses, in the return path, the width of the signal light pulses is widened again by the positively dispersed turbid medium and the imaging or processing distance adjuster, and then becomes short or ultra-short signal light pulses again by negative dispersion compression of the mirror negative dispersion generating device, and is received by the imaging receiver placed at the imaging position.

3. The apparatus for illuminating an inner optical layer for optical imaging and processing in a turbid medium according to claim 1, wherein the mirror-image negative dispersion generating means is mainly composed of a prism, a lens, etc., wherein an output surface of the dispersion compensating prism is formed by high-precision grinding and polishing controlled by a computer so as to satisfy a compensation requirement for optical path differences of two light beams adjacent to each pair of optical frequencies of the N light beams, that is, an absolute value of optical path differences of two light beams adjacent to each pair of optical frequencies generated by the negative dispersion generating means is equal to or approximately equal to an absolute value of optical path differences of two light beams having corresponding frequencies generated by an imaging or processing distance adjuster plus an optical path in the turbid medium, but both have opposite signs; and using an optical retroreflective sheeting to make the output surface of the dispersion compensating prism an optical retroreflective surface for retroreflecting the N light beams having different angles of incidence.

4. The device for optical imaging and processing of endo-optical layer illumination in turbid media according to claim 1, characterized in that said turbid media comprises human biological tissue, animal biological tissue, sea water, river water, lake water, pond water, fog, smog, haze, ice, snow, rain and any solid, liquid and gaseous medium with optically positive dispersion that is permeable to light waves but absorbing or/and scattering light waves.

5. The device for illuminating an inner optical layer for optical imaging and processing in a turbid medium according to claim 1, characterized in that the electromagnetic field frequency range of the N beams comprises the visible light region, or/and the infrared region, or/and the ultraviolet region, or/and the X-ray region.

6. The apparatus for endo-optical layer illumination for optical imaging and processing in turbid media according to claim 1, characterized in that said optical processing comprises various non-notched or notched medical lightwave treatments and treatments, underwater wireless optical communication, and various optical treatments involving the delivery of lightwave energy in turbid solid, liquid and gaseous media.

7. The apparatus for internal optical layer illumination for optical imaging and processing in a turbid medium according to claim 1, wherein the N optical waves of the internal optical layer illumination and processing apparatus are adapted to be generalized to the N acoustic waves, thereby producing an acoustic wave imaging and processing apparatus in a turbid medium, the apparatus comprising: an acoustic pulse generator; mirror negative dispersion generating means for acoustic waves; an acoustic imaging or machining distance adjuster; an acoustic imaging receiver.