KR20120082730A - Enhancement of retina imaging quality and microsurgery resolution by using optically filtered led light having maximum absorption by blood in retina vessels - Google Patents

Abstract

PURPOSE: A vascular imaging apparatus using a light emitting diode light source and a laser vascular surgery apparatus including the same are provided to improve surgery precision of retinal vessels by connecting the vascular imaging apparatus and a laser irradiation apparatus. CONSTITUTION: An LED system(110) generates an LED light source. A first optical system(120) includes a color filter changing the LED light source into an absorption wavelength range of blood inside a blood vessel. A monitoring apparatus(130) monitors a blood vessel tissue irradiating the LED light source having a wavelength range changed through the first optical system. The color filter is composed of a yellow filter(123) and a cyan filter(125) formed at the backend of the yellow filter.

Description

Vascular imaging apparatus using an LED light source and laser vascular surgical apparatus including the same {Enhancement of retina imaging quality and microsurgery resolution by using optically filtered LED light having maximum absorption by blood in retina vessels}

The present invention relates to a vascular imaging apparatus using an LED light source and a laser vascular surgery apparatus including the same, and more particularly, to the blood in the blood vessels in the retinal tissue by controlling the wavelength distribution of light generated from a white light emitting diode (LED). By maximizing the absorption of light by the vascular image relates to a vascular imaging device having a maximum contrast.

In addition, the present invention relates to a laser vascular surgical device that is excellent in the selective removal of the target site while minimizing damage to the retina by optimizing the distinction between surrounding blood vessels and tissues other than blood vessels.

The various internal structures forming the eye have very high transmittance in the near infrared (NIR) region, so that the near infrared ray can reach the inside of the eye very effectively. However, the application of sophisticated eye surgery is limited because the light of a typical NIR laser is not absorbed by certain retinal structures. This limitation can be solved when using ultrafast NIR laser sources. The focusing of NIR ultrafast laser pulses based on nonlinear absorption phenomena is illustrated by the clear example that allows removal of certain tissue sites with minimal incidental damage outside the focusing region. This technique makes it easy to realize the work of tissue coagulation by removal, drilling and delivery of laser irradiation within a small focal point. Ultra-fast laser-based surgery is an effective method for subcellular research, nervous system circulation, nervous system regeneration and nerve regeneration research, as it allows selective removal of ultra-fine targets with minimal damage to surrounding tissues. .

In the 1980s, many attempts were made to focus laser energy into the eye using a focused optical system with a numerical aperture. With this approach, limited removal of the retinal layer by Er: YAG laser at 2940 nm wavelength is the first selective intraocular tissue removal example. This approach, however, could not be applied to the removal of ILM without damaging the tissue of the lower retinal layer. The best solution for ex vivo retinal injury was 4 um with Er: YAG, but thermal degeneration was found in tissues even at 70 um depth from the time of removal.

Abnormal cell division and neovascularization represent symptoms of dangerous eye diseases such as proliferative diabetic retinopathy (PDR). Without proper treatment, at least 50% of patients with advanced PDR become blind within 5 years. In the treatment of PDR, thermal lasers are used to prevent retinal capillaries in which multiple hemorrhages occur (scatter photocoagulations) as a common treatment to slow the growth of new abnormal blood vessels. Treatment with this non-selective retinal injury method results in a permanent vision impairment. To overcome this damage, infrared laser sources, CO2 and Holmium YAG laser (via an optical fiber delivery system), are being piloted in eye disease surgery. However, incidental damage to surrounding tissues by shock wave effects and heat has been reported.

The phenomenon of tissue removal by laser can be described in optical breakdown or thermo-mechanical models. The optical breakdown model consists of plasma formation, shock wave formation, tissue removal and cavitation, and tissue damage. Plasma formation accounts for tissue removal for limited energy in materials with low laser light absorption or in a transparent state. The transient thermomechanical transient model covers all laser induced ablation reactions except photochemical laser ablation. In the absence of photochemical or state transitions, the entire energy absorbed by the material is used to increase the temperature of the object. In this state, the temperature of the object can be expressed as T (r) in any particular area (r), which is expressed as:

T (r) = r e (r) / Cv

Where r is the density of the material, e (r) is the absorbed energy density, and Cv is the specific heat capacity per unit volume. In addition, the redistribution of space by heat diffusion occurs when energy is absorbed in a particular area. This model explains that a limited occurrence of damage is observed if the pulse duration is less than the thermal relaxation time of the tissue volume.

Irradiation with a powerful ultrafast laser beam causes non-linear multiphoton absorption even if the target material does not have linear absorption in the laser wavelength region. The absorbed energy is propagated by energy relaxation via an oscillation (or phonon) mode with a time constant of several picoseconds in the adjacent region. On the other hand, when the pulse width of the laser is shorter than the time constant of the above propagation, it is mainly carried by electrons without thermal diffusion. Therefore, in the case of ultrafine surgery using ultrafast laser, thermal damage to surrounding tissues can be minimized and mechanical shock formation due to subsequent light induction to biological tissues is maintained unaffected. This effectively shows that the ultrafine laser surgery process is free of fever.

On the other hand, high density formation of free electrons generated through material excitation by laser focusing enables the formation of high density plasma in the target material. This is an artificially processable method of intracellular microstructures with no impact on cell survival, demonstrating the potential great potential for cell biology. Ultrafast lasers can also be applied to new forms of three-dimensional stereoscopic imaging techniques that produce high-resolution 3-D images by multi-photon absorption and emission based on nonlinear optical phenomena. The above technical advances can also take advantage of multiphoton absorption in photodynamic therapy. These lasers can emit almost exclusively light energy for retinal removal within the wavelength range of 700-1000 nm and are a potential source for biomedical applications for precise alteration in individual tissues.

Ultrafast lasers were already introduced at Interlase Inc. in 2001 for LASER-assisted in situ keratomileusis surgery. This shows that the only ultrafast lasers can be successfully applied to precise surgery on tissues with minimal damage to biological tissues.

The present invention has been made to solve the above problems, an object of the present invention, through the wavelength distribution control of the white light LED output to provide a high-quality image characteristic that maximizes the differentiation of blood vessels and other retinal tissue in the ocular fundus imaging In providing a device.

In addition, to provide a laser vascular surgery apparatus that effectively improves the vascular surgical accuracy on the retina by interlocking the blood vessel imaging device and a more controlled laser irradiation device.

The vessel imaging apparatus using the LED light source of the present invention comprises: an LED system 110 for generating an LED light source; A first optical system 120 having a color filter converting the LED light source into an absorption wavelength range of blood in a blood vessel; And a monitoring device 130 for monitoring blood vessel tissue irradiated with the LED light source whose wavelength range is converted through the first optical system 120. Characterized in that it comprises a.

In this case, the color filter may include a yellow filter 123 and a cyan filter 125 formed at a rear end of the yellow filter 123.

In addition, the LED light source is a white irradiation light is applied, the wavelength range of the LED light source converted through the color filter on the first optical system 120 is characterized in that 500nm ~ 700nm.

In addition, the blood vessel imaging apparatus is provided between the first optical system 120 and the blood vessel tissue, and the wavelength range of 540 ~ 550nm and 570 ~ 580nm for the LED light source whose wavelength range is converted through the first optical system 120 It is characterized in that it comprises a polarizing beam separator 140 for irradiating to the blood vessel tissue separated by.

In addition, a first polarized lens 151 is provided at the rear end of the cyan filter 125 on the first optical system 120, and a second polarized lens 152 is provided between the monitoring device 130 and the blood vessel tissue. It features.

In addition, the first optical system 120 includes a plurality of convex lenses 121 provided at the front end of the yellow filter 123 and a plurality of flat iron lenses provided between the yellow filter 123 and the cyan filter 125. It characterized in that it further comprises (124).

In addition, the blood vessel imaging device is characterized in that applied to the vascular tissue of the eye.

Laser vascular surgery apparatus of the present invention includes an LED system 110 for generating an LED light source; A first optical system 120 having a color filter converting the LED light source into an absorption wavelength range of blood in a blood vessel; A monitoring device (130) for monitoring blood vessel tissue irradiated with an LED light source whose wavelength range is converted through the first optical system (120); A laser light source 210 for generating a laser beam; A guide light source 220 for generating a guide beam; Fixing unit 230 for fixing the position of the eye to be treated; A first dichroic mirror 160 provided between the monitoring device 130 and the fixing part 230; A second optical system 240 for irradiating a laser beam and a guide beam generated by the laser light source and the guide light source to the fixed eye through the first dichroic mirror 160; And a galvano scanner 250 mounted on the second optical system 240 to control the direction of the laser beam and the guiding beam. Characterized in that it comprises a.

At this time, the pulse width of the laser beam is 150fs and the repetition rate is characterized in that 1kHz.

Finally, the diameter of the laser beam is characterized in that 2.0 ~ 2.4um when reaching the retina surface of the eye.

The vascular imaging apparatus of the present invention and the laser vascular surgery apparatus including the same according to the above-described configuration are selective of the retinal vascular wall without any damage to the surrounding retinal cell layer or the lower layer through an algorithm capable of analyzing retinal vascular images in real time. It has the effect of enabling perforation. These technological advances are expected to have applicability for early diagnosis of retinal diseases and ultra-precision therapy based on ultrafast lasers.

1 is a schematic view of a blood vessel imaging device of the present invention
2 is a light distribution diagram of absorption light and absorption wavelength of blood before and after white light LED wavelength distribution control.
3 is a schematic view of the vascular surgery apparatus of the present invention
4 is an intact intraretinal image obtained by changing the depth of focus based on the blood vessel imaging device of the present invention.
FIG. 5 is an image of FIG. 4 converted using a Prewitt operater.
6 is a graph showing contrast values obtained according to depths of focus while varying image analysis sizes.
FIG. 7 is a graph showing contrast values obtained according to depths of focus while varying image analysis positions.
8 is an intact intraocular retinal image obtained by varying the polarization angle of the polarizing plate of the irradiation light and the analysis light using the blood vessel imaging device of the present invention.
FIG. 9 is an image of FIG. 8 converted using a Prewitt operater.
10 is a graph showing the contrast values obtained by varying different image analysis sizes with respect to angles between polarizers.
FIG. 11 is a graph showing contrast values obtained by changing different image analysis positions with respect to angles between polarizers. FIG.
12 is a fundus image before ultrafast laser pulse irradiation for intraocular retinal ultrafine surgery.
Fig. 13 shows fundus image after ultrafast laser pulse irradiation for intraocular retinal ultrafine surgery.
14 is a digital camera image of the area undergoing surgery
FIG. 15 is a tomography of the surface of the retina and the tissues below the cell layer obtained after microsurgery of retinal tissue using ultrafast laser pulses of various fluences.
FIG. 16 shows three types of changes in vascular and retinal tissues: "no change", "removal of ILM only", and "vessel wall perforation" based on tomographic analysis of intact retinal tissue after a single pulse scan of ultrafast laser. Statistical processing of the correlation of damage by type

Hereinafter, an embodiment of a blood vessel imaging apparatus using the LED light source of the present invention as described above will be described in detail with reference to the accompanying drawings.

Referring to FIG. 1, a blood vessel imaging apparatus of the present invention includes a LED system 110 for generating an LED light source and a color filter for converting the LED light source into an absorption wavelength range of blood in a blood vessel. Monitoring device 130, polarizing beam splitter 140, the first and second polarizing lenses for monitoring the vascular tissue irradiated by the optical system 120 and the LED light source is converted wavelength range through the first optical system 120 151, 152, a first dichroic mirror 160, and an optometry lens 170.

The LED system 110 is composed of a white light LED (CREE XR-E Q5, 180lm) having a fast switching time. In addition, the LED system 110 is applied to the irradiation light to obtain a continuous fundus image in real time.

The first optical system 120 is configured to control the wavelength distribution of the LED light source generated from the LED system 110 and serves to irradiate the eye tissue with the LED light source.

The first optical system 120 includes a convex lens 121, an iris 122, a Y filter 123, a flat iron lens 124, a cyan filter 125, and a first mirror 126.

The convex lens 121 may be a lens having a convex centered convex lens. Through the convex lens 121 to focus the LED light source. The convex lens 121 is provided in plurality. The iris 122 is positioned on the focal point. The iris 122 controls the amount of light of the LED light source. The iris 122 may be applied to a conventional aperture configuration. The yellow filter 123 is positioned at the rear end of the iris 122. The yellow filter 123 is a normal yellow filter is applied, it is used to remove the sensitization of the rust oil and red oil emulsion. That is, the blue light of the white LED light source is removed. The flat iron lens 124 is positioned at the rear end of the yellow filter 123. The flat lens 124 refers to a lens having one flat surface and a convex other surface. In the present invention, the flat iron lens 124 is a combination of three flat iron lenses 124a, 124b, and 124c, and condenses the LED light source irradiated to the rear end of the iris 122. Done. First, second and third flat iron lenses are sequentially disposed from the rear end of the yellow filter 123. The first flat iron lens 124a has a focal length of 150 mm, the second flat iron lens 124b has a focal length of 300 mm, and the third flat iron lens 124c has a focal length of 200 mm. The cyan filter 125 is positioned at the rear end of the flat iron lens 124. The cyan filter 125 is a conventional cyan filter is applied, it is used to remove the redness of the rust emulsion. That is, the red light of the LED light source is removed. The LED light source whose wavelength distribution is controlled by the configuration as described above shows a wavelength distribution close to green light, and is irradiated to the polarization beam splitter 140 through the mirror 126.

The wavelength range of the LED light source controlled through the first optical system 120 is about 500 nm to 700 nm, as shown in FIG. 2.

In addition, the measured power of the LED light source meets international laser stabilization criteria, and as shown in FIG. 2, the wavelength of blood in the vascular tissue shows that there are two other maximum absorption bands corresponding to hemoglobin absorption at 542 nm and 576 nm. Able to know.

Therefore, the polarizing beam splitter 140 is configured at the rear end of the first optical system 120 of the present invention. The polarizing beam splitter 140 serves to irradiate the vascular tissue by separating the LED light source whose wavelength range is converted into the wavelength range of 540 ~ 550nm and 570 ~ 580nm through the first optical system 120.

The monitoring device 130 is divided into a CCD 131 and a control unit 132. The CCD 130 may be configured with a conventional charged couple device, and the controller 132 may be configured with a personal computer (PC) and a monitor.

In this case, the vascular imaging apparatus of the present invention has the following configuration so that light reflected directly from the cornea and lens of the eye tissue and eye fluid and the retina does not reach the CCD used as the photodetector.

A first polarization lens 151 is provided on the first optical system 120, and a second polarization lens 152 is provided between the monitoring device 130 and the polarization beam splitter 140. Therefore, light that is directly reflected through the first polarization lens 151 and the second polarization lens 152 does not reach the CCD 131.

The blood vessel imaging apparatus of the present invention also includes a first dichroic mirror filter 160. The first dichroic mirror 160 formally refers to a dichroic filter mirror. The first dichroic mirror 160, as a filter for color correction, has a characteristic of reflecting on the complementary color of a wavelength transmitted by the filter itself unlike other filters.

The first dichroic mirror 160 may be referred to as a Y (yellow), M (magenta), C (cyan) color correction filter installed in the color expander to adjust the color of the color print. In the case of a yellow filter, the filter itself looks like a blueish mirror. The mixing box may be used in combination to uniformize the LED light source to use the first dichroic mirror 160.

In other words, the dichroic filter may be a mirror that serves as a filter that reflects a range of wavelengths of visible light and transmits the remaining light. The first dichroic mirror 160 has a configuration for reflecting a laser beam of a vascular surgery apparatus, which will be described later.

Hereinafter, an embodiment of a laser vascular surgery apparatus including a blood vessel imaging apparatus using the LED light source of the present invention as described above will be described in detail with reference to the accompanying drawings.

Referring to FIG. 3, the vascular surgery apparatus of the present invention includes an LED system 110 for generating an LED light source and a first optical system including a color filter for converting the LED light source into an absorption wavelength range of blood in blood vessels. And a monitoring device 130 for monitoring blood vessel tissue irradiated with an LED light source whose wavelength range is converted through the first optical system 120, a laser light source 210 for generating a laser beam, and a guide beam. A guide light source 220 to be generated, a fixing part 230 for fixing a position of an eye to be treated, a first dichroic mirror 160 provided between the monitoring device 130 and the fixing part 230 and And a second optical system 240 for irradiating a laser beam and a guide beam generated by the laser light source and the guide light source to the fixed eye through the first dichroic mirror 160, and on the second optical system 240. Is installed on the laser And it consists of a galvanometer scanner (250) for controlling the direction of the guiding beam.

)레이저로 구성되며, 펄스폭은 150fs이고, 1kHz의 반복률을 갖는다. 상기 레이저 빔은 제2 미러(241), 제2 다이크로익 미러(244), 제3 미러(245)를 통해 제2 빔 파장기(246)에 전달되고, 최종적으로 갈바노 스캐너(250)에 전달된다.The laser light source 210 generates an ultrafast laser beam. The laser beam is Ti-sapphire (

It consists of a laser, pulse width is 150fs, and has a repetition rate of 1kHz. The laser beam is transmitted to the second beam wavelength 246 through the second mirror 241, the second dichroic mirror 244, and the third mirror 245, and finally to the galvano scanner 250. do.

)레이저로 구성된다. 상기 가이드 레이저 빔은 제 4 미러(242), 제5 미러(243), 제2 다이크로익 미러(244) 및 제3 미러(245)를 통해 제2 빔 파장기(246)에 전달되고, 최종적으로 갈바노 스캐너(250)에 전달된다.The guide light source 220 generates a guide laser beam. The guide laser beam is Melles Griot Gas laser (

It consists of a laser. The guide laser beam is transmitted to the second beam wavelength 246 through the fourth mirror 242, the fifth mirror 243, the second dichroic mirror 244, and the third mirror 245, and finally The galvano scanner 250 is delivered.

Accordingly, the laser beam, the guide laser beam, and the LED light source for the fundus image are combined with the optometry 170 through the first dichroic mirror 160. At this time, the laser beam and the guide laser beam passes through the galvano scanner 250 that can control the direction of the beam to apply the laser irradiation to the limited area of the eye. The optometry 170 has a focal length of 36.0 mm. At this time, the diameter of the laser beam is about 2.2um when reaching the retina surface. The eye to be tested is a pig eye in a chamber filled with 0.005M socruise solution (pH 7.4), and the cornea of the eye is fixed with a VFLATACS lens 232 to minimize the optical path change of the pig eye lens itself. Accordingly, the fixing part 230 is composed of an eyeball holder 231 and a VFLATACS lens 232.

Experimental Example 1

As shown in FIG. 4, in order to apply the imaging method of the present invention, an image of an intact porcine eye retina is first focused by using a color CCD having a resolution of 640 x 480 up and down about the image plane. Acquisition was made with a change of m to +140 m. As mentioned above, the wavelength distribution of the white light LED for illumination is controlled to be the maximum at the absorption wavelength of blood, so that the other tissues on the retina appear green, while the blood vessel part is weak, but the extra red wavelength diffuse reflection It looks red in color. This was chosen to obtain an image system for better basal image and real time monitoring during ultrafast laser irradiation of blood vessels.

FIG. 5 shows an image of the acquired image of FIG. 4 transformed using an algorithm based on the "Prewitt" operator described above. In particular, it should be noted that, as predicted, the interface and boundary between blood vessels and other retinal tissues are very clearly distinguished.

FIG. 6 shows a change in contrast depth for each contrast value by selecting a processing area having various sizes ranging from 25x25 pixels to the overall image size based on the center point of the acquired image.

FIG. 7 shows a result obtained by fixing the image size to be analyzed (50x50 pixels) and changing the selection area from the center to the edge of the image to be analyzed. From this, it can be seen that when the analysis image size is small in the central region of the fundus retinal image obtained, the method of the present invention is very sensitive to the change of imaging focus. LED light source has shown that it is possible.

Experimental Example 2

8 shows image acquisition results for polarization angles between inserted polarization control elements in order to minimize direct reflection from the eyeball construct. That is, the results of analyzing the quality of the retinal image while changing the polarization angle of the polarizing plate inserted to polarize the light of the illumination LED and the analytical polarizing plate located in front of the CCD detector. Optical designs using these orthogonal polarizers have already been adopted as a way to prevent direct reflective light from various intraocular tissues from reaching the detector CCD. The theoretical basis of this method has been reported to be a very effective method of acquiring better quality images in the double-pass technique, which was known half a century ago as a method of acquiring fundus photography. In the meantime, the above double-pass technique has been mainly applied as a method of acquiring images in the near infrared region. In the present invention, the above method was devised for the first time in the visible light region using a white light LED. Acquire images at various polarization angles at intervals of 0 degrees to 180 degrees and 5 degrees centering on the 90 degrees at which the polarization angles of the two polarizers are orthogonal under the imaging plane corresponding to the image focus (FIG. 8), and the "Prewitt operator" The converted image is shown in FIG. 9.

FIG. 10 illustrates a variation in polarization angle of each contrast value by selecting a processing area having various sizes ranging from a size of 25x25 pixels to an overall image size based on the center point of the acquired image.

FIG. 11 shows a result obtained by fixing the image size to be analyzed (50x50 pixels) and changing the selection area from the center to the edge of the image to be analyzed. From this, it can be seen that when the analysis image size is small in the central region of the fundus retinal image obtained by the present invention, the contrast values are very sensitive to the polarization angles of the two different polarizers. Imaging has been shown to be possible using white LED light sources with controlled wavelength distribution

Very fast (scanning time = 475 ms) laser scans a specific area on the retina of 275 x 250 mm size based on the retinal image obtained with an acquisition speed of about 25 frames / sec (fps) or more under the optimized optical system configuration. Could investigate. In this case, the number of laser scanning lines is 96, and the interval between consecutive scanning lines is 2.6 m. This scan condition means that there is no overlap between the ultrafast laser pulses, considering that the pulse train speed of the laser used is 1 kHz and the laser beam size is 2.2 mm in the focal plane. In particular, the dynamic image acquisition enables the pattern recognition process and the Galvano scanner used for laser delivery to track the area where the eye is initially set, even though it is hard to move.

Experimental Example 3

Figures 12 to 14 show fundus images before and after ultrafast laser pulse irradiation (before irradiation of Figure 12 and after irradiation of Figure 13) and visualized digital images (Cannon, Powershot A640) for intraocular retinal ultrafine surgery (Figure 14). ). It is seen in laser fluence of 50.6 J / cm2 or more that ultra-fast laser irradiation on the retinal surface containing primary blood vessels on the retina results in continuous ejection of blood due to destruction of the vessel wall. However, at low fluence (25.3 J / cm2), no blood ejection was observed.

Experimental Example 4

FIG. 15 shows a three-dimensional tomography of the surface of the retina and the tissues below the cell layer obtained after microretinal surgery using the ultrafast laser as shown in FIGS. 12 to 14. The tomography technique used was optical coherence tomography (OCT) to obtain continuous images of 50 mm intervals of the retinal cross section. In the OCT image acquisition, the refractive index is fixed at 1.4 and has a z-axis resolution of 4 mm. Meanwhile, the refractive index used was consistent with the water value according to the composition of the ocular vitreous body (96.4% of the composition is water). At laser fluence values lower than 2.5 J / cm2, both blood vessel and retinal surfaces were intact. The OCT image confirmed that the FLUUNUS, which can only damage blood vessels by laser scanning without any damage to the endothelial cell walls of blood vessels, was 5.1 J / cm2. The actual vascular perforation occurred at 25.3 J / cm2 with visible damage to the ILM covering the vessel wall. Laser fluence higher than 50.6 J / cm 2 was able to induce traces of bleeding from blood vessels by perfect perforation in the vessel wall. At increased laser fluence (101.2 J / cm 2), it can be seen that a significant amount of bleeding frequency in the blood vessels increases.

The most important observation is that even with increased laser fluence of up to 202.4 J / cm2, no visible damage to the layered structure supporting any lower retinal layer other than the superretinal ILM and vessel wall is observed. This result is quite different from previous similar experimental reports. Recent reports by Rockwell et al have reported that the intensity of ultrafast laser beams can lead to transmission through a nonlinear optical shape called laser self-focusing within the retinal cell layer, resulting in significant damage of the retinal cell layer. In the design of the present invention, the scan area (275 X 250 um) wider than the size of blood vessels was irradiated with accurate focal formation and a single pulse to the retina, which is a tissue around the retina other than the specific laser surgery caused by the low threshold of the blood vessel wall including hemoglobin. It is believed that no incidental damage from

Experimental Example 5

FIG. 16 shows OCT experimental results of intact porcine intraocular tissues by a single pulse scan of the ultrafast laser shown in FIG. 15. Three types of primary blood vessel change are "no change", "ILM only removal" and "vascular wall perforation." It shows the result of correlation statistical processing of each type of damages by classifying them as damage types. The removal of ILM on blood vessels is first observed at fluences of about 5.1 J / cm2, with an increase of 25.3 to 50.6 J / cm2 that appears more pronounced as the fluence increases and higher than 151.8 J / cm2 Is saturated. Vascular perforation, on the other hand, is shown by laser irradiation with a fluence higher than 12.7 J / cm 2. For determination of critical fluences for two different injuries, we estimate the percentage probability at low fluences. ILM damage critical fluence and vessel wall perforation on vessels were set at 2.0 J / cm2 and 5.3 J / cm2, respectively. It can be seen that this value is very consistent with the previously reported ex vivo state of ILM damage threshold of 2.2 J / cm2 in pig eye and 5.9 J / cm2 corresponding to vascular perforation.

The technical spirit should not be interpreted as being limited to the above embodiments of the present invention. Various modifications may be made at the level of those skilled in the art without departing from the spirit of the invention as claimed in the claims. Therefore, such improvements and modifications fall within the protection scope of the present invention, as will be apparent to those skilled in the art.

100: blood vessel imaging device
110: LED system 120: the first optical system
130: monitoring device 140: polarized beam splitter
151: first polarized lens 152: second polarized lens
160: first dichroic filter 170: optometry lens
200: blood vessel surgery apparatus
210: laser light source 220: guide light source
230: fixing part 240: second optical system
250: galvano scanner

Claims (10)

LED system 110 for generating an LED light source;
A first optical system 120 having a color filter converting the LED light source into an absorption wavelength range of blood in a blood vessel; And
A monitoring device (130) for monitoring blood vessel tissue irradiated with an LED light source whose wavelength range is converted through the first optical system (120);
Blood vessel imaging device using an LED light source comprising a.
The method of claim 1,
The color filter,
And a yellow filter (123) and a cyan filter (125) formed at a rear end of the yellow filter (123).
The method of claim 1,
White light is applied to the LED light source, and the wavelength range of the LED light source converted through the color filter on the first optical system 120 is 500nm to vascular imaging device using an LED light source, characterized in that.
The method of claim 3, wherein
The blood vessel imaging device,
It is provided between the first optical system 120 and the blood vessel tissue, and the LED light source is converted to a wavelength range of 540 ~ 550nm and 570 ~ 580nm through the first optical system 120 to irradiate the blood vessel tissue Blood vessel imaging apparatus using an LED light source, characterized in that it comprises a polarizing beam splitter (140).
The method of claim 2,
A first polarized lens 151 is provided at the rear end of the cyan filter 125 on the first optical system 120, and a second polarized lens 152 is provided between the monitoring device 130 and the blood vessel tissue. A blood vessel imaging device using an LED light source.
6. The method of claim 5,
The first optical system 120,
Further comprising a plurality of convex lenses 121 provided in front of the yellow filter 123 and a plurality of flat iron lens 124 provided between the yellow filter 123 and the cyan filter 125 Blood vessel imaging device using an LED light source.
The method of claim 1,
The blood vessel imaging device,
A blood vessel imaging device using an LED light source, characterized in that applied to the vascular tissue of the eye.
LED system 110 for generating an LED light source;
A first optical system 120 having a color filter converting the LED light source into an absorption wavelength range of blood in a blood vessel;
A monitoring device (130) for monitoring blood vessel tissue irradiated with an LED light source whose wavelength range is converted through the first optical system (120);
A laser light source 210 for generating a laser beam;
A guide light source 220 for generating a guide beam;
Fixing unit 230 for fixing the position of the eye to be treated;
A first dichroic mirror 160 provided between the monitoring device 130 and the fixing part 230;
A second optical system 240 for irradiating a laser beam and a guide beam generated by the laser light source and the guide light source to the fixed eye through the first dichroic mirror 160; And
A galvano scanner 250 mounted on the second optical system 240 and controlling the direction of the laser beam and the guiding beam;
Laser vascular surgery apparatus comprising a.
The method of claim 8,
And a pulse width of the laser beam is 150 fs and a repetition rate is 1 kHz.
The method of claim 8,
The diameter of the laser beam is a laser vascular surgery apparatus, characterized in that 2.0 ~ 2.4um when reaching the surface of the retina of the eye.

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