WO2020194712A1

WO2020194712A1 - Ophthalmic lens and method for producing ophthalmic lens

Info

Publication number: WO2020194712A1
Application number: PCT/JP2019/013802
Authority: WO
Inventors: 智裕川崎; 陽介奥平; 啓伊藤; 良一左高; 正宏水田
Original assignee: 株式会社ニコン
Priority date: 2019-03-28
Filing date: 2019-03-28
Publication date: 2020-10-01

Abstract

This ophthalmic lens is worn in or near an eyeball, wherein: the ophthalmic lens comprises a first region and a second region for transmitting light to form an image on the retina when the ophthalmic lens is worn on the eyeball; the first region intersects the optical axis of the ophthalmic lens and adds, to light rays passing through the first region, a first phase difference that monotonically increases or monotonically decreases in accordance with distance from the optical axis of the ophthalmic lens; the second region is located farther from the optical axis than the first region and adds, to light rays passing through the second region, a second phase difference that increases/decreases and continuously changes in accordance with distance from the optical axis; and the first phase difference and the second phase difference increase/decrease a total of 8 times or more in accordance with distance from the optical axis.

Description

Manufacturing method of ophthalmic lens and ophthalmic lens

The present invention relates to an ophthalmic lens such as an intraocular lens and a method for manufacturing an ophthalmic lens.

Intraocular lens (IOL) that is loaded into the eyeball instead of the crystalline lens after removal of the crystalline lens, intraocular lens (Phakic Intraocular lens, PIL) that is loaded into the eyeball in combination with the crystalline lens, contact lenses, etc. An intraocular lens that is used in contact with the eyeball of the lens is used.
In addition, a multifocal type ophthalmic lens is also used to supplement the focal adjustment ability of the eye. As one of the multifocal type ophthalmic lenses, a multifocal ophthalmic lens having a central refraction region that provides one refraction focusing ability and a diffraction region that provides near diffraction focusing ability and far diffraction focusing ability. Lenses have also been proposed (see Patent Document 1).

Japanese Patent Application Laid-Open No. 2009-82723

The ophthalmic lens of the first aspect of the present invention is an ophthalmic lens worn in or near the eyeball, and is a first region that transmits light and forms an image on the retina while being worn on the eyeball. And a second region, the first region intersects the optical axis of the ophthalmic lens and monotonically increases with respect to light rays passing through the first region according to the distance from the optical axis of the ophthalmic lens. Alternatively, a first phase difference that decreases monotonically is added, and the second region is located at a position farther from the optical axis than the first region, and is a distance from the optical axis with respect to a light ray passing through the second region. A second phase difference that increases or decreases according to the above and continuously changes is added, and the first phase difference and the second phase difference increase or decrease eight times or more in total according to the distance from the optical axis.
In the method for producing an ophthalmic lens according to the second aspect of the present invention, the ophthalmic lens according to the first aspect is produced by a processing apparatus.

It is a figure which shows the ophthalmic lens of the 1st Embodiment of this invention, FIG. 1 (a) shows the top view, and FIG. 1 (b) shows the sectional view. The figure which shows the cross section of the eyeball which attached the ophthalmic lens of 1st Embodiment. FIG. 3A is a diagram showing the relationship between the phase difference added to the aberration-free imaging ray by the ophthalmic lens of the first embodiment and the distance from the optical axis. FIG. 3B is a diagram showing an MTF of an image formed on the retina in an eyeball equipped with the ophthalmic lens of the first embodiment. FIG. 4A is a diagram showing the relationship between the phase difference added to the aberration-free imaging ray by the ophthalmic lens of Modification 1 and the distance from the optical axis. FIG. 4B is a diagram showing an MTF of an image formed on the retina in an eyeball equipped with the ophthalmic lens of the first modification. FIG. 5A is a diagram showing the relationship between the phase difference added to the aberration-free imaging ray by the ophthalmic lens of the second modification and the distance from the optical axis. FIG. 5B is a diagram showing an MTF of an image formed on the retina in an eyeball equipped with the ophthalmic lens of the second modification. FIG. 6A is a diagram showing the relationship between the phase difference added to the aberration-free imaging ray by the ophthalmic lens of Comparative Example 1 and the distance from the optical axis. FIG. 6B is a diagram showing an MTF of an image formed on the retina in an eyeball equipped with the ophthalmic lens of Comparative Example 1. FIG. 7A is a diagram showing the relationship between the phase difference added to the aberration-free imaging ray by the ophthalmic lens of the modified example 3 and the distance from the optical axis. FIG. 7B is a diagram showing an MTF of an image formed on the retina when the pupil diameter is 3 mm in the eyeball to which the ophthalmic lens of the modified example 3 is attached. FIG. 7 (c) is a diagram showing an MTF of an image formed on the retina when the pupil diameter is 4 mm, similar to FIG. 7 (b). FIG. 8A is a diagram showing the relationship between the phase difference added to the aberration-free imaging ray by the ophthalmic lens of the modified example 4 and the distance from the optical axis. FIG. 8B is a diagram showing an MTF of an image formed on the retina when the pupil diameter is 3 mm in the eyeball to which the ophthalmic lens of the modified example 4 is attached. FIG. 8 (c) is a diagram showing an MTF of an image formed on the retina when the pupil diameter is 4 mm, which is the same as in FIG. 8 (b). FIG. 9A is a diagram showing the relationship between the phase difference added to the aberration-free imaging ray by the ophthalmic lens of the modified example 5 and the distance from the optical axis. FIG. 9B is a diagram showing an MTF of an image formed on the retina when the pupil diameter is 3 mm in the eyeball to which the ophthalmic lens of the modified example 5 is attached. FIG. 9C is a diagram showing an MTF of an image formed on the retina when the pupil diameter is 4 mm, which is the same as in FIG. 9B. FIG. 10A is a diagram showing the relationship between the phase difference added to the aberration-free imaging ray by the ophthalmic lens of the modified example 6 and the distance from the optical axis. FIG. 10B is a diagram showing an MTF of an image formed on the retina when the pupil diameter is 3 mm in the eyeball to which the ophthalmic lens of the modified example 6 is attached. FIG. 10 (c) is a diagram showing an MTF of an image formed on the retina when the pupil diameter is 4 mm, similar to FIG. 10 (b). FIG. 11A is a diagram showing an MTF of an image formed on the retina in an eyeball equipped with a conventional single focus type ophthalmic lens. FIG. 11B is a diagram showing an MTF of an image formed on the retina in an eyeball equipped with a conventional double focus type ophthalmic lens. FIG. 12 is a diagram showing an example of the function fγ (x) of the equation (2).

(First Embodiment)
FIG. 1 is a diagram showing an intraocular lens 10 as an example of an ophthalmic lens according to an embodiment of the present invention, FIG. 1 (a) is a top view, and FIG. 1 (b) is FIG. 1 (a). The cross-sectional view on the Y axis is shown. The intraocular lens 10 has two refracting surfaces, a light incident surface 11 and a light emitting surface 12. As an example, the incident surface 11 is a spherical surface, and the injection surface 12 is a surface having smooth irregularities with respect to the reference surface 13 which is a spherical surface.

The optical axis AX of the intraocular lens 10 is the axis of rotational symmetry of the entrance surface 11 and the emission surface 12 (or the reference surface 13). The radius r0 of the intraocular lens 10 is, for example, about 2.5 mm to 3.0 mm. The intraocular lens 10 may have a support portion (not shown) around the intraocular lens 10.
The X-axis and the Y-axis shown in FIG. 1A are axes in an arbitrary direction that are in a plane perpendicular to the optical axis AX and are orthogonal to each other. For any one point Q on the injection surface 12, the distance from the optical axis is defined as the distance r, and the azimuth angle from the X axis centered on the optical axis AX is θ.

FIG. 2 is a diagram showing a cross section of the eyeball 100 loaded with the intraocular lens 10 of the first embodiment. The intraocular lens 10 shown in FIG. 2 is a lens loaded in the eyeball 100 in place of a crystalline lens (not shown) extracted from the eyeball, and is a vitreous body originally at a position where the crystalline lens is arranged. It is arranged between 32 and the anterior chamber 31. An iris 36 is arranged inside the anterior chamber 31. The intraocular lens 10 is arranged so that its optical axis AX substantially coincides with the center line EX of the eyeball on a straight line connecting the center of the cornea 30 and the center of the macula 35, but it does not necessarily have to coincide exactly. ..

The intraocular lens 10 is a lens having the same refractive power as the original crystalline lens, and the light rays L1 and L2 from an object in the outside world (not shown) are refracted by the cornea 30, and the anterior chamber 31, the intraocular lens 10, And through the lens 32, an image is formed at an imaging point 34 near the center of the yellow spot 35 on the retina 33.
The two rays L1 and L2 shown in FIG. 2 are examples, and in reality, a large number of rays (imaging rays) emitted from an object and entering the incident surface are the above-mentioned cornea 30, the intraocular lens 10, and the like. The image is formed at the imaging point 34 on the retina 33. If there is no aberration as a whole in the optical system including the cornea 30, the anterior chamber 31, the intraocular lens 10, and the vitreous 32, there is no optical path length difference (phase difference) between each of the large number of imaging rays. Hereinafter, the aberration-free imaging ray is referred to as an aberration-free imaging ray.

As an example, the amount of smooth unevenness of the injection surface 12 shown in FIG. 1B with respect to the reference surface 13 is determined according to the distance from the optical axis AX in the direction away from the optical axis AX, as will be described later. Therefore, the uneven shape described above is, for example, a shape rotationally symmetric with respect to the optical axis AX. In FIG. 1B, the uneven shape of the injection surface 12 is exaggerated in the direction of the optical axis AX.
In the present specification, the distance from the reference surface 13 to the injection surface 12 in each portion of the injection surface 12 is referred to as the height H of the injection surface 12, and the side where the injection surface 12 is separated from the intraocular lens 10 is a positive reference numeral. Expressed as.

The refractive index of the intraocular lens 10 is similar to the refractive power of the crystalline lens, which is 1.494 as an example, while the refractive index of the vitreous body in contact with the intraocular lens 10 is 1.336 as an example. Therefore, for the light rays passing through the portion where the height of the ejection surface 12 is H, the optical path length difference of H × (1.494-1.336) = H × 0.158 with respect to the light rays passing through the portion where the height is 0. Is added. The phase difference of light is (2π / wavelength) times the optical path length difference.

Therefore, if the wavelength of the passing light beam is 0.55 [μm], the light ray passing through the portion of the ejection surface 12 having a height of H [μm] passes through the portion having a height of 0 [μm]. A phase difference of 2π × H × 0.158 / 0.55 = H × 1.805 [rad] is added to the light beam.
Therefore, the height H [μm] of the injection surface 12 from the reference surface 13 is equivalent to the phase difference [rad] obtained by multiplying this value by 1.805 as an example. Therefore, in the present specification, the phase differences P1 and P2 [rad] in which the intraocular lens 10 adds the height H [μm] of the ejection surface 12 from the reference surface 13 to the aberration-free imaging ray. Treat as equivalent to.

FIG. 3A shows the relationship between the phase differences P1 and P2 added to the aberration-free imaging light by the ejection surface 12 of the ophthalmic lens 10 of the first embodiment and the distance r from the optical axis AX. It is a figure.
In the first embodiment, as shown in FIGS. 1 (b) and 3 (a), the height H of the injection surface 12 is higher than the reference surface 13 in the region near the optical axis AX. There is. Therefore, the phase difference P1 has a positive value in the region near the optical axis AX.

The phase difference P1 decreases monotonically as the distance r from the optical axis AX increases, and the first minimum value MP1 is set at a position where the distance r from the optical axis AX becomes r1 (about 0.6 mm as an example). Take. Then, in the region where the distance r from the optical axis AX is r1 or more, the phase difference P2 repeats increasing and decreasing according to the distance r from the optical axis AX.

Hereinafter, the region in which the distance r from the optical axis AX is within r1 is referred to as the first region Z1, and the region in which the distance r from the optical axis AX is r1 or more is referred to as the second region Z2. That is, the first region Z1 adds a first phase difference P1 that monotonically decreases with respect to the light rays passing through the first region Z1 according to the distance r from the optical axis AX, and the second region Z2 provides the second region Z2. A second phase difference P2 that is increased or decreased a plurality of times according to the distance r from the optical axis AX and is continuous is added to the passing light beam.

The difference between the maximum value and the minimum value of the first phase difference P1 is called the first fluctuation width W1, and the difference between the maximum value and the minimum value of the second phase difference P2 is called the second fluctuation width W2.
Further, the first phase difference P1 and the second phase difference P2 are also collectively referred to as phase differences P1 and P2.

In FIG. 1A, the optical axis AX intersects the injection surface 12 in the first region Z1, that is, the first region Z1 intersects the optical axis AX.
In the first embodiment, the first phase difference P1 and the second phase difference P2 are continuous at the boundary BL between the first region Z1 and the first region Z2, and have a minimum value MP1.

The light rays passing through the first region Z1 and the second region Z2 form an image on the macula 35 on the retina 33 by the refraction action of the cornea 30, the intraocular lens 10, the vitreous body 32 and the like.
In the intraocular lens 10 of the first embodiment, the aberration-free imaging ray when the first phase difference P1 and the second phase difference P2 are 0 produces an image of an object at infinity without aberration from the macula 35 to 0.21 [mm. ] Is set to form an image at a position distant from the cornea 30 side. This is done by optimizing the shapes and arrangements of the incident surface 11 and the reference surface 13 of the intraocular lens 10 by a known method. In other words, when the height H of the ejection surface 12 is 0 over the entire surface of the first region Z1 and the second region Z2, the eyeball 100 loaded with the intraocular lens 10 has eyeballs 100 to 1.79. [m] An image of a distant object is imaged on the macula 35 without aberration.

In the first embodiment, the phase differences P1 and P2 added to the light rays transmitted through the intraocular lens 10 by the height H of the ejection surface 12 are cos with respect to the distance r from the optical axis represented by the equation (1). It is an amount according to the function Pa (r) represented by the function.

... (1)

Here, fγ in the equation (1) is a function represented by the following equation (2).

... (2)
In equation (2), sgn (x) is a function that returns the sign of the argument x.

If 2πr ² / r _s ² is set as x, fγ in the equation (1) is expressed as fγ (cos (x)).
FIG. 12 is a graph showing the function fγ (cos (x)) at γ = 1.5 and γ = 3 together with cos (x). When the constant γ increases from 1, the function fγ (cos (x)) increases its slope in the vicinity of the 0 cross point (x = π / 2 or 3π / 2) of cos (x).
On the other hand, although not shown, when the constant γ decreases from 1, the function fγ (cos (x)) has an inclination in the vicinity of the 0 cross point (x = π / 2 or 3π / 2) of cos (x). It will be relaxed.

In the definition of the equation (2), the function fγ (cos (x)) cannot be defined because it includes division by zero at γ = 1. However, when γ = 1, the function fγ (cos (x)) may be considered to match cos (x).
Returning to equation (1), the constant Ps is a constant representing the magnitude of the component of the function fγ (cos (x)) using the cos function for the square of the distance r from the optical axis.

In the phase differences P1 and P2 [rad] shown in FIG. 3 (a), the values of the constants of the equations (1) and (2) are r _s = 1.053 [mm], P _s = 1.1 [rad], and It is a function Pa (r) when γ = 2.03.
As can be seen from equation (1), phase difference P1, P2 in accordance with the period determined by the constant r _s, increases or decreases according to the change in the distance r from the optical axis AX. In the first embodiment, since the constant r _s is 1.053 [mm], while the distance r changes from r = 0 on the optical axis AX to r = 3 [mm] on the outermost circumference of the intraocular lens 10. , The phase differences P1 and P2 are increased or decreased eight times in total.

The increase / decrease in the phase differences P1 and P2 accompanying the change in the distance r adds a function as a multifocal filter to the transmitted light transmitted through the first region Z1 and the second region Z2. As a result, the intraocular lens 10 can increase the depth of focus of the optical system including the cornea 30, the anterior chamber 31, the intraocular lens 10, and the vitreous body 32.

FIG. 3B shows an imaging point 34 in the eyeball 100 (an optical system including the cornea 30, the anterior chamber 31, the intraocular lens 10, and the vitreous body 32) loaded with the intraocular lens 10 of the first embodiment. A graph showing the simulation results at each defocus position (Defocus Position) of an MTF (Modulation Transfer Function) at a predetermined spatial frequency (eg, about 50 [LP / mm]) of an image formed on a nearby retina 33. Is.

The simulation is disclosed in Chapter 36, Section 4 of "Handbook of Optical Systems: Vol. 4 Survey of Optical Instruments" (WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim) compiled by Herbert Gross in 2008. , Based on the Navarro model.
That is, assuming an optical system with a focal length based on the Navarro model (an optical system including a cornea 30, anterior chamber 31, intraocular lens 10, and vitreous body 32), each imaging ray passing through the optical system is shown in the figure. The simulation was performed by adding the phase differences P1 and P2 shown in 3 (a).

However, as described above, the intraocular lens 10 of the first embodiment adds a function as a multifocal filter for transmitted transmitted light. Therefore, in the simulation, the -1st order diffracted light (diffracted light to which a negative refractive power is added) from the multifocal filter formed by the first region Z1 and the first region Z2 is on the horizontal axis of FIG. 3 (b). The condition of the intraocular lens 10 is set so that the defocus position is focused at the position of 0 [mm]. Specifically, this is done by finely adjusting the position of the intraocular lens 10 and the shapes of the incident surface 11 and the reference surface 13. This also applies to the simulation of the intraocular lens 10 such as each modification described later.

FIG. 3B shows the MTF when the pupil diameter (diameter of the opening of the iris 36) is 3 mm. The pupil diameter of the human eyeball 100 is about 3 to 4 mm in a bright environment and about 6 mm in a dark environment.
The vertical axis of the graph represents the modulation transfer rate (MTF) of the spatial frequency component (eg, about 50 [LP / mm]) of the image on the retina 33, and the horizontal axis represents the amount of defocus in the center line EX direction (eg, about 50 [LP / mm]). The retina 33 is set as the reference = 0, and the side closer to the cornea 30 is negative) is expressed in [mm] units.
The length of 0.36 [mm] on the horizontal axis corresponds to the refractive power of 1 diopter. Therefore, for example, when an image of an object at infinity is formed at a position of 0 [mm] on the horizontal axis, an image of an object 1 m away from the eyeball 100 is located at a position of -0.36 [mm] on the horizontal axis. It is formed. This also applies to the simulation of the intraocular lens 10 such as each modification described later.

As described above, the intraocular lens 10 of the first embodiment functions as a multifocal filter for transmitted light transmitted through the intraocular lens 10. That is, the intraocular lens 10 of the first embodiment divides a plurality of light rays (luminous flux) passing through the first region Z1 and the second region Z2 into a plurality of light fluxes. Then, with the 0th-order diffracted light as a reference (0 diopter), a refractive power (addition) of −0.75 diopters is added to the -1st-order diffracted light among the plurality of light beams divided into wave planes, and +0 to the +1st-order diffracted light. The refractive power (addition) of .75 diopters is added. The refractive power of 0.75 diopters corresponds to a relative position of about 0.27 [mm] as the defocus position shown on the horizontal axis of FIG. 3 (b).

On the other hand, FIG. 11A is a diagram showing an MTF when a conventional single focus type intraocular lens is attached to the eyeball as a comparative example. The pupil diameter used in the simulation and the number of spatial frequencies of the image on the retina 33 are the same as in FIG. 3 (b). Note that FIG. 11A shows the simulation results in the two cases of the pupil diameter of 3 mm and the pupil diameter of 4 mm.

Comparing FIG. 3 (b) and FIG. 11 (a), the intraocular lens 10 of the present embodiment is relatively high in a wide range of about −0.6 [mm] to 0 [mm] as the defocus position. It can be seen that the value of MTF can be obtained. The range of this defocus position corresponds to the range of the distance from the eyeball 100 from 50 [cm] to infinity as the distance to the object, and the object at a wide distance is placed on the retina 33 with relatively high contrast. Can be imaged. That is, the intraocular lens 10 of the first embodiment increases the depth of focus of the optical system including the cornea 30, the anterior chamber 31, the intraocular lens 10, and the vitreous 32.

FIG. 11B is a diagram showing an MTF when a conventional bifocal intraocular lens is attached to the eyeball as a comparative example. The pupil diameter and the number of spatial frequencies of the image on the retina 33 used in the simulation are the same as in the case of FIG. 11A.
Even in the conventional double focus type intraocular lens, the MTF value can be increased especially at the defocus position 0 [mm] and the defocus position -0.8 [mm], but for example, the defocus position −0. The value of MTF at 4 [mm] is low.

On the other hand, the intraocular lens 10 of the present embodiment has a relatively high MTF in a wide range of about −0.6 [mm] to 0 [mm] as the defocus position, as shown in FIG. 3 (b). It turns out that the value of can be obtained.
Since the pupil diameter of 3 mm is generally the pupil diameter in a bright place such as outdoors in the daytime, the intraocular lens 10 of the present embodiment has an infinite distance from the eyeball 100 from 50 [cm] in a bright place. A good image can be obtained for a wide range of objects up to a long distance.

(Modification example 1)
FIG. 4A is a diagram showing the relationship between the phase differences P1 and P2 added to the aberration-free imaging light beam by the ejection surface 12 of the intraocular lens 10 of the modification 1 and the distance r from the optical axis AX. Is. Since the configuration of the intraocular lens 10 of the first modification is almost the same as the configuration of the intraocular lens 10 of the first embodiment described above, only the differences from the first embodiment described above will be described below. ..
In the phase differences P1 and P2 [rad] shown in FIG. 4 (a), the values of the constants of the equations (1) and (2) are r _s = 0.900 [mm], P _s = 1.1 [rad], and It is a function Pa (r) when γ = 2.03.

In the first modification, since the constant r _s is 0.900 [mm], while the distance r changes from r = 0 on the optical axis AX to r = 3 [mm] on the outermost circumference of the intraocular lens 10. The phase differences P1 and P2 are increased or decreased 11 times in total. As the number of times of this increase / decrease increases, the refractive power (addition) applied to each diffracted light transmitted by the first region Z1 and the second region Z2 as the multifocal filter also increases from the case of the above-described first embodiment. Become stronger. In the first modification, the refractive power (addition) of -1 diopter is added to the -1st order diffracted light among the plurality of light beams divided on the wave surface with the 0th order diffracted light as a reference (0 diopter), and the +1 diopter is added to the +1 diopter. Is added with a refractive power (addition) of +1 diopter. The refractive power of 1 diopter corresponds to a relative position of about 0.36 [mm] as the defocus position shown on the horizontal axis of FIG. 4 (b).

FIG. 4B shows the defocus of the MTF at a predetermined spatial frequency of the image formed on the retina 33 near the imaging point 34 in the eyeball 100 loaded with the intraocular lens 10 of the modification 1. It is a graph which shows the simulation result at a position. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

As shown in FIG. 4B, the intraocular lens 10 of the modified example 1 has an image MTF at each defocus position of 0 [mm], -0.36 [mm], and -0.72 [mm] and its vicinity. It can be increased, that is, the depth of focus can be increased.
This is because phase differences P1 and P2 that increase or decrease 11 times according to the distance r from the optical axis AX are added to the light rays that pass through the first region Z1 and the second region Z2, and the phase differences P1 and P2 are multiplexed. This is because it functions as a focus filter.

(Modification 2)
FIG. 5A is a diagram showing the relationship between the phase differences P1 and P2 added to the aberration-free imaging light beam by the ejection surface 12 of the intraocular lens 10 of the modification 2 and the distance r from the optical axis AX. Is. Since the configuration of the intraocular lens 10 of the second modification is almost the same as the configuration of the intraocular lens 10 of the first embodiment described above, only the differences from the first embodiment described above will be described below. ..
In the phase differences P1 and P2 [rad] shown in FIG. 5 (a), the values of the constants of the equations (1) and (2) are r _s = 0.747 [mm], P _s = 1.1 [rad], and It is a function Pa (r) when γ = 2.03.

In the second modification, since the constant r _s is 0.747 [mm], while the distance r changes from r = 0 on the optical axis AX to r = 3 [mm] on the outermost circumference of the intraocular lens 10. The phase differences P1 and P2 are increased or decreased 16 times in total. As the number of times of this increase / decrease increases, the refractive power (addition) applied to each diffracted light transmitted by the first region Z1 and the second region Z2 as the multifocal filter also increases in the above-described first embodiment and modified examples. It becomes stronger than the case of 1. In the second modification, the refractive power (addition) of -1.5 dioptres is added to the -1st-order diffracted light among the plurality of light beams divided on the wave surface with the 0th-order diffracted light as a reference (0 diopter), and the +1 next time A refractive power (addition) of +1.5 diopters is added to the folding light. The refractive power of 1 diopter corresponds to a relative position of about 0.50 [mm] as the defocus position shown on the horizontal axis of FIG. 4 (b).

FIG. 5B shows each defocus of the MTF at a predetermined spatial frequency of the image formed on the retina 33 near the imaging point 34 in the eyeball 100 loaded with the intraocular lens 10 of the second modification. It is a graph which shows the simulation result at a position. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

As shown in FIG. 5 (b), the intraocular lens 10 of the modified example 2 has an MTF of the image at each defocus position of 0 [mm], -0.50 [mm], and -1 [mm] and its vicinity. Can be enhanced. In each of the above defocus positions, the distance from the eyeball 100 corresponds to infinity, 66 [cm], and 33 [cm], respectively, as the distance to the object. That is, the depth of focus of the eyeball 100 can be increased by the intraocular lens 10 of the modification 2.
This is because phase differences P1 and P2 that increase or decrease 16 times according to the distance r from the optical axis AX are added to the light rays that pass through the first region Z1 and the second region Z2, and the phase differences P1 and P2 are multiplexed. This is because it functions as a focus filter.

(Comparative Example 1)
FIG. 6A is a diagram showing the relationship between the phase differences P1 and P2 added to the aberration-free imaging light by the ejection surface 12 of the intraocular lens 10 of Comparative Example 1 and the distance r from the optical axis AX. Is. The configuration of the intraocular lens 10 of Comparative Example 1 is almost the same as the configuration of the intraocular lens 10 of the first embodiment described above, but differs from the first embodiment described above in the following points.
In the phase differences P1 and P2 [rad] shown in FIG. 6 (a), the values of the constants of the equations (1) and (2) are r _s = 1.255 [mm], P _s = 1.1 [rad], and It is a function Pa (r) when γ = 2.03.

That is, the value of the constant r _s is the 1.255 [mm], it is greater than the value of the constant r _s in the first embodiment and the modifications described above. As a result, in the intraocular lens 10 of Comparative Example 1, the phase difference P1 and P2 are changed while the distance r changes from r = 0 on the optical axis AX to r = 3 [mm] on the outermost circumference of the intraocular lens 10. Increases or decreases only 5 times in total. As a result, the refractive power (addition) applied to each of the diffracted light transmitted by the first region Z1 and the second region Z2 as the multifocal filter is also weaker than in the case of the first embodiment and each modification described above. ..

In Comparative Example 1, the refractive power (addition) of −0.5 dioptre is added to the -1st-order diffracted light among the plurality of light beams divided on the wave surface with the 0th-order diffracted light as a reference (0 diopter), and the +1 next time. A refractive power (addition) of +0.5 diopters is added to the folding light. The refractive power of 0.5 diopters corresponds to a relative position of about 0.18 [mm] as the defocus position shown on the horizontal axis of FIG. 6 (b).

FIG. 6B shows the MTFs of the images formed on the retina 33 near the imaging point 34 in the eyeball 100 loaded with the intraocular lens 10 of Comparative Example 1 at a predetermined spatial frequency. It is a graph which shows the simulation result at a focus position. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

As shown in FIG. 6B, the intraocular lens 10 of Comparative Example 1 can increase the MTF of the image at the defocus position of 0 to −0.36 [mm]. However, the range of this defocus position corresponds to the range in which the distance from the eyeball 100 is from 1 [m] to infinity as the distance to the object. Therefore, the intraocular lens 10 of Comparative Example 1 cannot improve the contrast of the image with respect to an object at a short distance of less than 1 [m] from the eyeball 100.
This is because the intraocular lens 10 of Comparative Example 1 has phase differences P1 and P2 added to the light rays passing through the first region Z1 and the second region Z2 only five times according to the distance r from the optical axis AX. This is because it does not increase or decrease, and therefore the refractive power (addition) applied as a multifocal filter is weak.

(Supplementary information about the first embodiment, the first modification, and the second modification)
On the other hand, in the intraocular lens 10 of the first embodiment and each modification described above, the phase differences P1 and P2 added to the light rays passing through the first region Z1 and the second region Z2 are from the optical axis AX. Increase or decrease 8 times or more in total according to the distance r.
With this configuration, it is possible to give an appropriate multiplex effect to the light rays transmitted through the first region Z1 and the second region Z2, and it is possible to realize an ophthalmic lens 10 having an increased depth of focus.

There is no particular upper limit to the number of times the phase differences P1 and P2 increase or decrease according to the distance r from the optical axis AX, but if the number of times of increase or decrease is too large, the first region Z1 and the second region Z2 as the multifocal filter The refractive power (addition) added to the transmitted light becomes too strong. Therefore, in reality, the refractive power (addition) applied to the ± 1st-order diffracted light by the first region Z1 and the second region Z2 may be up to about ± 2 diopters. In this case, the eyeball 100 provided with the intraocular lens 10 has an improved MTF for an image of an object from a distance of 25 [cm] from the eyeball 100 (+4 diopters) to infinity (0 diopters). The degree of addition of ± 2 diopters corresponds to 22 times as the number of times of increase / decrease of the phase differences P1 and P2 with the change of the distance r from the optical axis AX.

Further, in the intraocular lens 10 of the modified example 1 and the modified example 2, the second region Z2 is added to the transmitted light beam in the range where the distance r from the optical axis AX is 1.5 mm or more and 3 mm or less. The phase difference P2 increases or decreases 7 times or more according to the distance r from the optical axis AX.
With this configuration, an appropriate multiplex effect can be given by the light rays transmitted through the first region Z1 and the second region Z2, and the ophthalmic lens 10 having a further increased depth of focus can be realized.

(Modification 3)
In the intraocular lens 10 of the above-mentioned

modifications

1 and 2, a relatively large refractive power (with respect to the -1st-order diffracted light and the + 1st-order diffracted light transmitted through the first region Z1 and the second region Z2 and divided into wave planes ( Since the addition degree) is added, the depth of focus can be greatly increased.
However, since the refractive power (addition) applied to the -1st-order diffracted light and the + 1st-order diffracted light is large, the MTFs shown in FIGS. 4 (b) and 5 (b) are at specific discrete defocus positions. Is high, but it is a little low in the middle.
The intraocular lens 10 of the third modification improves the MTF value not only in the above-mentioned specific discrete defocus positions but also in the middle of these positions.

FIG. 7A is a diagram showing the relationship between the phase differences P1 and P2 added to the aberration-free imaging light beam by the ejection surface 12 of the intraocular lens 10 of the modification 3 and the distance r from the optical axis AX. Is. Since the configuration of the intraocular lens 10 of the modification 2 is almost the same as the configuration of the intraocular lens 10 of the modification 1 described above, only the differences from the modification 1 described above will be described below.

The phase differences P1 and P2 that the ejection surface 12 of the intraocular lens 10 of the modified example 3 shown in FIG. 7A adds to the aberration-free imaging light beam are the functions Pb (r) of the following equation (3). It is represented by.

... (3)
The first term on the right side of equation (3) is the argument of the cos function, which is the argument of the function fγ on the right side of equation (1), plus an arbitrary initial phase b. The second term of the right side of formula (3) is a Gaussian function with respect to the distance r from the optical axis AX, is a constant P _g is a constant representing the magnitude of the components of the Gaussian function, constant r _g is the Gaussian function Is a constant that represents the width of the distribution of.

In the phase differences P1 and P2 [rad] of the intraocular lens 10 of the modified example 3 shown in FIG. 7 (a), the values of the constants of the above equations (3) and (2) are r _s = 0.900 [ _{mm], P s = 1.1 [} rad], γ = 2.03, b = 0 [rad], r g = 0.24 [mm], is equal to the value of the function Pb (r) in the case of _{P g} = 3.40 [rad].
That is, the phase differences P1 and P2 of the intraocular lens 10 of the modified example 7 are the Gaussian of the second term of the equation (3) in the phase differences P1 and P2 of the ejection surface 12 of the intraocular lens 10 of the modified example 1 described above. It is the addition of a function.

In the third modification, since the constant r _s is 0.900 [mm] as in the first modification, the distance r is from r = 0 on the optical axis AX to r = 3 [mm] on the outermost circumference of the intraocular lens 10. ], The phase differences P1 and P2 increase or decrease 11 times in total. Therefore, as in the above-described modification 1, in the first region Z1 and the second region Z2, the refractive power (addition) of -1 diopter is added to the -1st-order diffracted light with reference to the 0th-order diffracted light, and +1. A refractive power (addition) of +1 diopter is added to the next diffracted light.

In Modification 3, the phase difference P1, P2, to the extent of a distance r from the optical axis AX up to 2 times the constant r _g is the width of the Gaussian distribution, the second term of equation (3) Gauss The component by the function is added. However, in the phase differences P1 and P2 shown in FIG. 7A, the distance r1 from the optical axis AX to the first minimum value MP1 which is the boundary BL between the first region Z1 and the second region Z2 is 0.7 [mm]. and is a distance of about three times the constant r _{g =} 0.24 [mm]. Therefore, the increase of the phase differences P1 and P2 by the Gaussian function of the second term of the equation (3) is substantially limited to the range of the first region Z1 closer to the optical axis AX than the above distance r1. There is.

As a result, in the modified example 3, the second fluctuation width W2 of the second phase difference P2 in the second region Z2 is twice the above constant P _s , but the phase difference P1 in the first region Z1. The first fluctuation width W1 is obtained by adding the above constant P _g to the fluctuation width W2. That is, the first fluctuation width W1 is larger than the second fluctuation width W2 by a constant P _g = 3.40 [rad].

7 (b) and 7 (c) show the image formed on the retina 33 near the imaging point 34 at a predetermined spatial frequency in the eyeball 100 loaded with the intraocular lens 10 of the modification 3. It is a graph which shows the simulation result at each defocus position of MTF. FIG. 7B shows the simulation results for each case with a pupil diameter of 3 mm, and FIG. 7C shows the simulation results for each case with a pupil diameter of 4 mm. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

As shown in the simulation result when the pupil diameter is 3 mm shown in FIG. 7 (b), the intraocular lens 10 of the modified example 3 is 0 [, similar to the intraocular lens 10 of the modified example 1 shown in FIG. 4 (b). At each defocus position of -0.36 [mm], -0.72 [mm] and its vicinity, the MTF of the image can be increased, that is, the depth of focus can be increased.

Further, the intraocular lens 10 of the modified example 3 is an intermediate position of each of the above defocus positions as compared with the intraocular lens 10 of the modified example 1 shown in FIG. 4 (b), which is -0.18 [mm],-. It can be seen that the MTF of the image can be relatively increased even at each defocus position of 0.55 [mm].
Further, in the simulation result when the pupil diameter is 4 mm shown in FIG. 7 (c), the MTF of the image is enhanced at each defocus position of -0.18 [mm] and -0.55 [mm]. I understand.

(Modification example 4)
FIG. 8A is a diagram showing the relationship between the phase differences P1 and P2 added to the aberration-free imaging light beam by the ejection surface 12 of the intraocular lens 10 of the modified example 4 and the distance r from the optical axis AX. Is. Since the configuration of the intraocular lens 10 of the modification 4 is almost the same as the configuration of the intraocular lens 10 of the modification 3 described above, only the differences from the modification 3 described above will be described below.

The phase differences P1 and P2 [rad] of the intraocular lens 10 of the modified example 4 shown in FIG. 8 (a) are constants of the constants of the above equations (3) and (2) in the modified example 3. Only P _g is equal to the value of the function Pb (r) changed to 5.0 [rad].
That is, in the intraocular lens 10 of the modified example 4, the first fluctuation width W1 of the first phase difference P1 is larger than the second fluctuation width W2 of the second phase difference P2 by a constant P _g = 5.0 [rad].

8 (b) and 8 (c) show the image formed on the retina 33 near the imaging point 34 at a predetermined spatial frequency in the eyeball 100 loaded with the intraocular lens 10 of the modification 4. It is a graph which shows the simulation result at each defocus position of MTF. FIG. 8B shows the simulation results for each case with a pupil diameter of 3 mm, and FIG. 8C shows the simulation results for each case with a pupil diameter of 4 mm. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

From the simulation results shown in FIGS. 8 (b) and 8 (c), the intraocular lens 10 of the modified example 4 had a relatively low MTF in the intraocular lens 10 of the above-mentioned modified example 1 -0.18 [mm. ], It can be seen that the MTF can be increased even at each defocus position of -0.55 [mm].

(Modification 5)
FIG. 9A is a diagram showing the relationship between the phase differences P1 and P2 added to the aberration-free imaging light beam by the ejection surface 12 of the intraocular lens 10 of the modified example 5 and the distance r from the optical axis AX. Is. Since the configuration of the intraocular lens 10 of the modification 5 is almost the same as the configuration of the intraocular lens 10 of the modification 3 described above, only the differences from the modification 3 described above will be described below.

The phase differences P1 and P2 [rad] of the intraocular lens 10 of the modified example 5 shown in FIG. 9 (a) are constants of the constants of the above equations (3) and (2) in the modified example 3. Only P _g is equal to the value of the function Pb (r) changed to 12.5 [rad].
That is, in the intraocular lens 10 of the modified example 5, the first fluctuation width W1 of the first phase difference P1 is larger than the second fluctuation width W2 of the second phase difference P2 by a constant P _g = 12.5 [rad].

9 (b) and 9 (c) show the image formed on the retina 33 near the imaging point 34 at a predetermined spatial frequency in the eyeball 100 loaded with the intraocular lens 10 of the modified example 5. It is a graph which shows the simulation result at each defocus position of MTF. FIG. 9B shows the simulation results for each case with a pupil diameter of 3 mm, and FIG. 9C shows the simulation results for each case with a pupil diameter of 4 mm. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

From the simulation results shown in FIGS. 9 (b) and 9 (c), the intraocular lens 10 of the modified example 5 had a relatively low MTF in the intraocular lens 10 of the above-mentioned modified example 1 -0.18 [mm. ], It can be seen that the MTF can be increased even at each defocus position of -0.55 [mm].

(Modification 6)
FIG. 10A is a diagram showing the relationship between the phase differences P1 and P2 added to the aberration-free imaging light beam by the ejection surface 12 of the intraocular lens 10 of the modified example 6 and the distance r from the optical axis AX. Is. Since the configuration of the intraocular lens 10 of the modification 6 is almost the same as the configuration of the intraocular lens 10 of the modification 3 described above, only the differences from the modification 3 described above will be described below.

The phase differences P1 and P2 [rad] of the intraocular lens 10 of the modified example 6 shown in FIG. 10 (a) are constants of the constants of the above equations (3) and (2) in the modified example 3. Only P _g is equal to the value of the function Pb (r) changed to 1.55 [rad].
That is, in the intraocular lens 10 of Comparative Example 2, the first fluctuation width W1 of the first phase difference P1 is larger than the second fluctuation width W2 of the second phase difference P2 by a constant P _g = 1.55 [rad].

10 (b) and 10 (c) show the image formed on the retina 33 near the imaging point 34 at a predetermined spatial frequency in the eyeball 100 loaded with the intraocular lens 10 of the modification 6. It is a graph which shows the simulation result at each defocus position of MTF. FIG. 10B shows the simulation results for each case with a pupil diameter of 3 mm, and FIG. 10C shows the simulation results for each case with a pupil diameter of 4 mm. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

In the simulation result when the pupil diameter is 4 mm shown in FIG. 10 (c), even in the intraocular lens 10 of the modified example 6, it is relatively high at each defocus position of -0.18 [mm] and -0.55 [mm]. It turns out that MTF can be obtained.
However, in the simulation results when the pupil diameter is 3 mm shown in FIG. 10 (b), no improvement in MTF is observed at each defocus position of -0.18 [mm] and -0.55 [mm].
In the modified example 6, the first fluctuation width W1 of the first phase difference P1 is increased by 1.55 [rad] from the second fluctuation width W2 of the second phase difference P2. However, with the increase of 1.55 [rad], a sufficient depth of focus expansion effect due to the first phase difference P1 could not be obtained as compared with the above-mentioned modified examples 3 to 5 in which the increase was increased by 2 [rad] or more.

(Supplementary explanation of modification 3, modification 4, and modification 5)
On the other hand, in the modified example 3, the modified example 4, and the modified example 5, the first fluctuation width W1 of the first phase difference P1 is 2 [rad] or more than the second fluctuation width W2 of the second phase difference P2. It is increasing. In other words, the first region Z1 has a position of 2 [rad] or more with respect to a light ray passing through at least a part of the first region Z1 (for example, a light ray passing through the optical axis AX) as compared with a light ray passing through the second region Z2. A phase difference is added.

Therefore, the intraocular lenses 10 of the modified examples 3, the modified examples 4, and the modified examples 5 have a multifocal effect due to the light rays transmitted through the second region Z2 and a depth of focus expanding effect due to the light rays transmitted through the first region Z1. Together, the depth of focus can be further increased.
There is no upper limit to the difference between the first fluctuation width W1 and the second fluctuation width W2, but if the difference is too large, the first fluctuation width W1 becomes too large, and the light rays transmitted through the first region Z1 become flares. It may be scattered in the field of view and reduce the contrast of the image. In order to prevent this, the difference between the first fluctuation width W1 and the second fluctuation width W2 can be set to about 50 [rad] or less.

(Supplementary explanation for the first embodiment and each modification)
In the above first embodiment and each modification, when the constant P _s of the equation (1) or the equation (3) is increased, the first fluctuation width W1 of the first phase difference P1 and the second phase difference P2 2 The fluctuation width W2 increases, respectively. As a result, the amount of light of the -1st-order diffracted light and the + 1st-order diffracted light diffracted from the first region Z1 and the second region Z2 as the multiple focus filter increases with respect to the light amount of the 0th-order diffracted light. Therefore, the multiplex effect of the intraocular lens 10 can be increased or decreased by increasing or decreasing the value of the constant P _s (increasing or decreasing the first fluctuation width W1 and the second fluctuation width W2).

When the value of γ of the function fγ (cos (x)) included in the equations (1) and (3) is 1, that is, when it is considered that fγ (cos (x)) = cos (x), the constant Ps When the value of is about 1.45 [rad], the amounts of the -1st-order diffracted light, the 0th-order diffracted light, and the + 1st-order diffracted light generated from the first region Z1 and the second region Z2 are substantially equal. Therefore, by setting the value of the constant Ps to about 0.75 to 1.75 [rad], the first region Z1 and the second region Z2 can function as an appropriate multiple focus filter.

In the above first embodiment and each modification, cos (2πr ² / r _s ² ) in the formula (1) may be expressed as cos (ar ² ) by substituting 2π / r _s ² with a. it can. Similarly, cos (2πr ² / r _s ^2- b) in the equation (3) can also be expressed as cos (ar ^2- b).
Further, as described above, the function fγ in the equations (1) and (3) is a function that increases the inclination of the cos function, which is an argument, near the 0 cross point. Therefore, fγ (cos (ar ² )) can also be expressed by, for example, the equation (4).

... (4)
Here, j is an arbitrary natural number, m is a natural number larger than about 3, and g _j is an arbitrary number.

^{Further, fγ (cos (ar 2 -b} )) , for example, it may be represented by the formula (5).

... (5)
Therefore, in the second phase difference P2 in the first embodiment and each modification, the distance r from the optical axis AX, an arbitrary constant a, and an arbitrary constant b are used, and j in the equation (5) is 0. of, it can also be interpreted to include components represented by cos ^{ar 2 -b}.

The second phase difference P2 in the first embodiment and the modified examples, the formula (5) where j is 1, also be interpreted to include components represented by cos ^{3ar 2 -3b} it can.
The term of the square of the distance r (r ² ) included in the argument of the cos function in the equations (1), (3), (4), and (5) is not necessarily the square of the distance r. It does not have to be, and it may be any of the distance r from the 1.5th power to the 2.5th power.

The value of the constant a in the above equations (4) and (5) is a parameter that determines the amount corresponding to the period of the cos function included in the equations (4) and (5), and therefore, the first region. It is a parameter that determines the refractive power of the above-mentioned multifocal filter formed by Z1 and the second region Z2.
The second term of the equation (3) is not limited to the Gaussian function described above, and may be a sinc function with respect to the distance r from the optical axis or a function represented by the “power power” of the sinc function.

Further, the phase difference P1 in the first region Z1 may be a function based on the above-mentioned Gaussian function or sin function plus another function. That is, the first phase difference P1 may include a component represented by a Gaussian function or a sinc function.
The shape of the height H is designed by making the shape of the height H of the injection surface 12 in the first phase difference P1, that is, the first region Z1 a shape based on the Gaussian function and the sinc function that are mathematically easy to handle. And easy to manufacture.

The amount of the phase differences P1 and P2 added to the transmitted light rays by the first region Z1 and the second region Z2 is a function of the above equations (1), (3), (4), and (5). The shape is not limited to the above shape, and other shapes may be used.
In any case, the phase differences P1 and P2 may be increased or decreased twice or more in the range from the optical axis AX to 1.5 mm according to the distance from the optical axis AX. As a result, an appropriate multifocal effect can be given to the transmitted light rays, and the depth of focus can be further increased.

Further, the second fluctuation width W2 of the second phase difference P2 is 1.5 [rad] or more and 3.5 [rad] or less in the range where the distance r from the optical axis AX is 1 mm to 1.5 mm (described above). The range of the value of the constant Ps may be twice as large as 0.75 to 1.75 [rad]). When the distance r is in the above range and the second fluctuation width W2 of the second phase difference P2 is in the above range, the -1st order diffracted light, the 0th order light diffracted light, and the +1st order diffracted light generated from the second region Z2 are set. The balance of the amount of light can be set to an appropriate value, and the multifocal effect can be exhibited more effectively.

In the above-described first embodiment and each modification, the amount of the second phase difference P2 added to the transmitted light ray by the second region Z2 is mainly cos with respect to the square of the distance r from the optical axis AX. It changes in proportion to (ar ² ). Therefore, the cycle of increase / decrease of the phase differences P1 and P2 according to the distance r from the optical axis AX is substantially inversely proportional to the distance r from the optical axis AX. This relationship is similar to the period of the lens height of the so-called Blaze type diffractive lens.

When the period of increase / decrease of the second phase difference P2 is substantially inversely proportional to the distance r from the optical axis AX, the -1st-order diffracted light and the + 1st-order diffracted light generated by dividing the wave plane by the second region Z2 have the distance r from the optical axis AX. Approximately equal refractive power is applied regardless of. As a result, the -1st-order diffracted light and the + 1st-order diffracted light can be focused at predetermined focus positions (defocus positions), respectively.

However, the cycle of increase / decrease of the second phase difference P2 does not have to be substantially inversely proportional to the distance r from the optical axis AX. In this case, different refractive powers are applied to the -1st-order diffracted light and the + 1st-order diffracted light generated by dividing the wave plane by the first region Z1 and the second region Z2 according to the distance r from the optical axis AX. As a result, aberrations corresponding to so-called spherical aberrations are added to the -1st-order diffracted light and the + 1st-order diffracted light, and each of them is focused in a range having a width in the center line EX direction from a predetermined focus position. As a result, the effect of further increasing the depth of focus can be obtained.
The cycle of increase / decrease of the second phase difference P2 may be monotonically increasing or monotonically decreasing with respect to the distance r from the optical axis AX.
Alternatively, the cycle of increase / decrease of the second phase difference P2 is constant and does not have to change with respect to the distance r from the optical axis AX.

The shape of the incident surface 11 and the reference surface 13 of the intraocular lens 10 may be a so-called aspherical surface such as an ellipsoidal surface, an eccentric surface, a hyperboloid or a paraboloid, instead of a spherical surface, and may be a concave surface or a convex surface. It may be flat or flat.
Further, the uneven shape may be formed on the incident surface 11 instead of the injection surface 12, or may be formed on both the incident surface 11 and the injection surface 12.

The intraocular lens 10 is made of, for example, an acrylic resin material (for example, a copolymer of acrylate and methacrylate), hydrogel, or silicone. Further, the ophthalmic lens (eg, intraocular lens 10) may be a flexible lens made of a foldable flexible material (eg, acrylic resin material, silicone), and is a lens made of a hard material. You may.
The refractive index of the material constituting the intraocular lens 10 is not limited to 1.494 described above, and may be any other value.

When the pupil diameter is 3 mm, in order to obtain both the multiple focus effect due to the first phase difference P1 and the second phase difference P2 and the depth of focus expansion effect due to the first phase difference P1, the pupil diameter is 3 mm. It is necessary that the first region Z1 and the second region exist in an appropriate area ratio inside.
Therefore, the distance r1 from the optical axis AX at the boundary BL between the first region Z1 and the second region Z2 may be 0.4 mm or more and 0.8 mm or less. If the distance r1 is smaller than 0.4 mm, the area of the first region Z1 becomes smaller, and a sufficient depth of focus expansion effect may not be obtained. On the other hand, when the distance r1 is larger than 0.8 mm, the area of the second region Z2 becomes small, and a sufficient multiplex effect may not be obtained.

In the above-described first embodiment and the intraocular lens 10 of each modification, the height H (phase difference P1, P2) of the ejection surface 12 is determined according to the distance r from the optical axis AX, that is, with respect to the optical axis AX. It is assumed that it has a rotationally symmetric shape. However, the phase differences P1 and P2 do not necessarily have to be rotationally symmetric with respect to the optical axis AX in order to correct astigmatism caused by the shape of the cornea 30.

For example, the phase differences P1 and P2 may have a shape that is twice symmetrical with respect to the optical axis AX, or may have a shape that is three times symmetrical. In this case, in order to determine the phase differences P1 and P2 according to the equations (1) and (2), for example, the value of r in the equations (1) and (2) is shown in the equation (6). The height P2 may be determined by replacing it with r'.
r'= r × {1 + g · sin (2θ－φ)} ・・・ (6)
Here, g is a constant of about 0 or more and 0.5 or less, θ is the azimuth angle of any one point Q on the injection surface 12 as described above, and φ is an arbitrary initial phase.

In the above first embodiment and each modification, the first phase difference P1 added to the transmitted light by the first region Z1 is assumed to decrease monotonically according to the distance r from the optical axis. However, the phase difference P1 may increase monotonically according to the distance r from the optical axis.
For example, the phase differences P1 and P2 shown in FIG. 3A and the like may have their symbols inverted (inverted in the vertical axis direction).
In that case, the boundary BL between the first region Z1 and the second region Z2 is the position where the phase differences P1 and P2, which increase monotonically according to the distance r from the optical axis, take the first maximum value. Become.

Further, in the first embodiment and each modification, the first phase difference P1 and the second phase difference P2 are assumed to be continuous at the boundary BL. However, the first phase difference P1 and the second phase difference P2 may change discontinuously at the boundary BL.
However, since the first phase difference P1 and the second phase difference P2 are continuous at the boundary BL, the scattered light generated by a sudden change in the phase difference is reduced, and the intraocular lens 10 provides a clearer field of view. Can be realized.

Further, in the first embodiment and each modification, the first phase difference P1 added to the transmitted light by the first region Z1 is assumed to continuously change according to the distance r from the optical axis. However, the phase difference P1 may change discretely according to the distance r from the optical axis.
However, since the first phase difference P1 continuously changes according to the distance r, the scattered light generated by the sudden change in the phase difference is reduced, and the intraocular lens 10 that provides a clearer field of view is realized. be able to.

Further, in both the first embodiment and each modification, the phase differences P1 and P2 are not only continuous but also smoothly continuous at the boundary BL between the first region Z1 and the second region Z2. good. In this case, it is possible to realize an intraocular lens 10 that reduces scattered light generated by a sudden change in the phase difference near the boundary BL and provides a clearer field of view.

In the above-described first embodiment and each modification, the phase differences P1 and P2 are added by changing the height H of the incident surface 11 or the ejection surface 12 of the intraocular lens 10. The method of adding the phase differences P1 and P2 is not limited to this method.
For example, when the intraocular lens 10 is composed of a plurality of lenses, a phase difference may be added by changing the shape (height) of the internal surface on which the lenses face each other.

Alternatively, the above phase difference can be added by providing a refractive index fluctuation portion in the first region Z1 and the second region Z2 of the intraocular lens 10 whose refractive index changes with respect to the distance r from the optical axis AX. can do. That is, for example, the intraocular lens 10 can be formed by using a material (eg, silicone, acrylic resin, etc.) whose refractive index changes concentrically around the optical axis AX.

Since the shape and refractive index of the crystalline lens disclosed in Table 36-16 of the above-mentioned document referred to in the above simulation are only standard examples, the refractive power of the actual intraocular lens 10 varies from person to person. Strictly speaking, it is different from the refractive power of.
Therefore, in the simulation using the Navarro model, the thickness of the intraocular lens 10 and the radius of curvature of the incident surface 11 and the reference surface 13 are appropriately changed so that the imaging point 34 (focus) coincides with the retina 33 of the Navarro model. Then perform the simulation.

Conversely, using the simulation method described above, it is also possible to determine whether or not any intraocular lens corresponds to one of the intraocular lenses 10 of the first embodiment or each modification. .. That is, the refractive index and shape of the intraocular lens are numerically attached to the Navarro model, and the radius of curvature of the incident surface 11 and the reference surface 13 of the intraocular lens is appropriately changed to form the imaging point 34 (focus). Is matched on the retina 33 of the Navarro model. Under that condition, it can be determined whether or not the phase differences P1 and P2 of the light rays passing through the first region Z1 and the second region Z2 have the above phase difference.

The ophthalmic lens in the present embodiment is not limited to the intraocular lens 10 (IOL) loaded in the eyeball 100 instead of the crystalline lens, but is an implantable contact lens (IPL) loaded between the iris 36 and the crystalline lens. ) May be. Further, the intraocular lens for a so-called piggy bag, which is additionally loaded for correction with respect to the eyeball 100 to which the intraocular lens is attached, may be used. Alternatively, it may be a corneal inlay or a corneal inlay loaded into the cornea.

Alternatively, the ophthalmic lens may be a contact lens worn on the outside of the cornea 30. In this case, the user can wear a contact lens having the configuration described in the present embodiment and an existing intraocular lens (eg, a single focus type IOL) on the eye and use them in combination. In addition, ophthalmic lenses can be used for various vision correction applications such as IOLs that can be used for both pseudo-lens and crystalline lenses.
Furthermore, the ophthalmic lens may be a spectacle lens worn away from the eyeball.

(Effects of the first embodiment and each modification)
(1) The ophthalmic lens (intraocular lens) 10 of the above first embodiment and each modification is an ophthalmic lens 10 worn in or near the eyeball, and emits light while being attached to the eyeball. It includes a first region Z1 and a second region Z2 that transmit and form an image on the retina 33. Then, the first region intersects the optical axis AX of the ophthalmic lens 10 and monotonically increases or decreases with respect to the light rays passing through the first region Z1 according to the distance r from the optical axis AX of the ophthalmic lens 10. The first phase difference P1 is added, and the second region Z2 is located at a position farther from the optical axis AX than the first region Z1 and is at a distance r from the optical axis AX with respect to a light ray passing through the second region Z2. A second phase difference P2 that increases or decreases accordingly and changes continuously is added, and the first phase difference P1 and the second phase difference P2 increase or decrease eight times or more in total according to the distance r from the optical axis AX. ..
With this configuration, it is possible to add an appropriate multiplex effect to the light rays transmitted through the first region Z1 and the second region Z2, and it is possible to realize an ophthalmic lens 10 having an increased depth of focus.

(2) The first phase difference P1 and the second phase difference P2 increase or decrease twice or more according to the distance r from the optical axis AX in the range where the distance r from the optical axis AX is up to 1.5 mm. Therefore, it is possible to add a more appropriate multifocal effect to the light rays transmitted through the first region Z1 and the second region Z2.
(3) The first phase difference P1 and the second phase difference P2 are continuous at the boundary BL between the first region Z1 and the second region Z2, and have a minimum value or a maximum value at the boundary BL. This makes it possible to realize an intraocular lens 10 that reduces scattered light generated by a sudden change in phase difference and provides a clearer field of view.

(4) By setting the distance from the optical axis AX to the boundary BL to 0.4 mm or more and 0.8 mm or less, the multifocal effect due to the first phase difference P1 and the second phase difference P2 and the first phase difference P1 The effect of expanding the depth of focus can be exhibited in a well-balanced manner.
(5) The first region Z1 is configured to add a phase difference of 2 [rad] or more to the light rays passing through at least a part of the first region Z1 as compared with the light rays passing through the second region Z2. The depth of focus can be further increased by the combination of the multifocal effect of the light rays transmitted through the second region Z2 and the depth of focus expansion effect of the light rays transmitted through the first region Z1.

(Manufacturing method of ophthalmic lenses)
Next, a method for manufacturing the ophthalmic lens 10 of the present embodiment described above will be described. The method for manufacturing the ophthalmic lens 10 is the method for manufacturing the ophthalmic lens described in the above-described first embodiment and each modification, and the ophthalmic lens is manufactured by using design data indicating the lens shape of the ophthalmic lens. It is provided with a processing process manufactured by a processing apparatus (eg, mold processing apparatus, cutting apparatus, polishing apparatus, etc.). The method for manufacturing an ophthalmic lens includes a design process for designing the lens shape and generating the design data. Here, the above design data can be generated by converting information (eg, design conditions) such as the above phase differences P1 and P2 into a lens shape (surface shape) and a refractive index distribution.

It is also possible to provide a lens set by providing a plurality of ophthalmic lenses of the present embodiment. In this case, the lens set includes a plurality of ophthalmic lenses having different depths of focus and the number of focal points (eg, multiple focal points such as two or three focal points) of the ophthalmic lens of the present embodiment.
The present invention is not limited to the above contents. Other aspects considered within the scope of the technical idea of the present invention are also included within the scope of the present invention. In this embodiment, all or a part of the above-described embodiments may be combined.

10: Ophthalmic lens (intraocular lens), 100: Eyeball, 11: Incident surface, 12: Ejection surface, 13: Reference surface, AX: Optical axis, r: Distance from optical axis, Z1; First region, Z2 2nd region, BL: boundary, P1: 1st phase difference: P2: 2nd phase difference, W1: 1st fluctuation width, W2: 2nd fluctuation width, 30: cornea, 31: anterior chamber, 32: vitreous body , 33: Retina, 35: Yellow spot, 36: Iris

Claims

In an ophthalmic lens worn in or near the eyeball
It has a first region and a second region that allow light to pass through and form an image on the retina when worn on the eyeball.
The first region intersects the optical axis of the ophthalmic lens and monotonically increases or decreases with respect to the light rays passing through the first region according to the distance from the optical axis of the ophthalmic lens. Add phase difference,
The second region is located at a position farther from the optical axis than the first region, and with respect to light rays passing through the second region, it increases or decreases depending on the distance from the optical axis and continuously changes. Add a second phase difference,
An ophthalmic lens in which the first phase difference and the second phase difference increase or decrease eight times or more in accordance with the distance from the optical axis.
In the ophthalmic lens according to claim 1.
An ophthalmic lens in which the first phase difference and the second phase difference increase or decrease twice or more according to the distance from the optical axis within a range of a distance from the optical axis up to 1.5 mm.
In the ophthalmic lens according to claim 1 or 2.
An ophthalmic lens in which the first phase difference and the second phase difference are continuous at a boundary between the first region and the second region, and have a minimum value or a maximum value at the boundary.
In the ophthalmic lens according to claim 3.
An ophthalmic lens in which the distance from the optical axis to the boundary is 0.4 mm or more and 0.8 mm or less.
In the ophthalmic lens according to any one of claims 1 to 4.
An ophthalmic lens in which the cycle of the increase / decrease of the second phase difference in each part of the second region differs depending on the distance from the optical axis.
In the ophthalmic lens according to any one of claims 1 to 5.
An ophthalmic lens in which the cycle of the increase / decrease of the second phase difference in each part of the second region increases or decreases monotonically according to the distance from the optical axis.
In the ophthalmic lens according to any one of claims 1 to 5.
An ophthalmic lens in which the cycle of the increase / decrease of the second phase difference in each part of the second region is substantially inversely proportional to the distance from the optical axis.
In the ophthalmic lens according to any one of claims 1 to 7.
The first region is an ophthalmic lens that adds a phase difference of 2 [rad] or more to a light ray passing through at least a part of the first region as compared with a light ray passing through the second region.
In the ophthalmic lens according to any one of claims 1 to 8.
The fluctuation range of the increase / decrease of the second phase difference is 1.5 [rad] or more and 3.5 [rad] or less in the range of the distance from the optical axis of 1 mm to 1.5 mm for ophthalmology. lens.
In the ophthalmic lens according to any one of claims 1 to 9.
The first phase difference is an ophthalmic lens containing a component represented by a Gaussian function with respect to the distance from the optical axis.
In the ophthalmic lens according to any one of claims 1 to 9.
The first phase difference is an ophthalmic lens containing a component represented by a function based on a sinc function with respect to the distance from the optical axis.
In the ophthalmic lens according to any one of claims 1 to 11.
The second phase difference is cos {ar (2 + β) using a distance r from the optical axis, an arbitrary constant a, an arbitrary constant b, and a real number β of −0.5 or more and 0.5 or less. An ophthalmic lens containing a component represented by −b}.
In the ophthalmic lens according to claim 12.
The second phase difference is an ophthalmic lens containing a component represented by cos {3ar (2 + β) -3b}.
The ophthalmic lens according to any one of claims 1 to 13.
An ophthalmic lens in which the first phase difference and the second phase difference increase or decrease 11 times or more in accordance with the distance from the optical axis.
The ophthalmic lens according to any one of claims 1 to 14.
In the second region, the light flux passing through the second region is divided into at least two light fluxes.
An ophthalmic lens that adds a degree of addition of 1 diopter or more between the two light beams.
The ophthalmic lens according to any one of claims 1 to 15.
The second phase difference added by the second region is an ophthalmic lens that increases or decreases seven times or more in a range of 1.5 mm or more and 3 mm or less from the optical axis.
The ophthalmic lens according to any one of claims 1 to 16.
An ophthalmic lens comprising a surface having an uneven shape that adds the first phase difference and the second phase difference to the light beam.
The ophthalmic lens according to any one of claims 1 to 17.
An ophthalmic lens in which the first phase difference and the second phase difference are rotationally symmetric with respect to the optical axis.
The ophthalmic lens according to any one of claims 1 to 17.
An ophthalmic lens in which the first phase difference and the second phase difference are two or more times symmetrical with respect to the optical axis.
In the ophthalmic lens according to any one of claims 1 to 19.
The ophthalmic lens is an intraocular lens that is loaded into the eyeball instead of the crystalline lens.
In the ophthalmic lens according to any one of claims 1 to 19.
The ophthalmic lens is an ophthalmic lens which is an implantable contact lens loaded between the iris and the crystalline lens.
In the ophthalmic lens according to any one of claims 1 to 19.
The ophthalmic lens is an ophthalmic lens which is a corneal inlay or a corneal inlay loaded into the cornea.
In the ophthalmic lens according to any one of claims 1 to 19.
The ophthalmic lens is an ophthalmic lens which is a contact lens that comes into contact with the cornea.
A method for manufacturing ophthalmic lenses
A method for manufacturing an ophthalmic lens, wherein the ophthalmic lens according to any one of claims 1 to 23 is manufactured by a processing apparatus.