WO2020194713A1 - Ophthalmic lens and ophthalmic lens production method - Google Patents

Abstract

This ophthalmic lens is placed within or near an eye, and is provided with a first region and a second region which transmit light and form an image on the retina when the lens is placed within/near the eye. The first region adds, to light rays passing through the first region, a first phase difference which monotonically increases or monotonically decreases according to the distance from the optical axis of the ophthalmic lens. The second region adds, to light rays passing through the second region, a second phase difference which increases or decreases multiple times according to the distance from the optical axis. The first phase difference varies in an amplitude which is 1.1 times or more that of the second phase difference.

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 in a state of being worn on the eyeball, transmits light to form an image on the retina. The first region includes a first region and a second region, and the first region adds a first phase difference that monotonically increases or decreases monotonically depending on the distance from the optical axis of the ophthalmic lens with respect to the light beam passing through the first region. Then, the second region adds a second phase difference that increases or decreases a plurality of times according to the distance from the optical axis to the light beam passing through the second region, and the fluctuation range of the first phase difference is the said. An ophthalmic lens that is 1.1 times or more the fluctuation range of the second phase difference.
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 Modification 3 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 Modification 3. 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 4 and the distance from the optical axis. FIG. 7B is a diagram showing an MTF of an image formed on the retina in the eyeball equipped with the ophthalmic lens of the modified example 4. 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 5 and the distance from the optical axis. FIG. 8B is a diagram showing an MTF of an image formed on the retina in an eyeball equipped with the ophthalmic lens of the modified example 5. FIG. 9A is a top view of the ophthalmic lens 10a of the modified example 6. FIG. 9B is a top view of the ophthalmic lens 10b of the modified example 7. 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 8 and the distance from the optical axis. FIG. 10B is a diagram showing an MTF of an image formed on the retina in an eyeball equipped with the ophthalmic lens of the modified example 8. The figure which shows the MTF of the image formed on the retina in the eyeball which attached the conventional single focus type ophthalmic lens.

(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. In this case, for example, the injection surface 12 may be a continuous 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 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 of the injection surface 12, and the side of the injection surface 12 away from the intraocular lens 10 is represented by a positive sign. ..

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 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 ophthalmic lens 10 of the first embodiment and the distance from the optical axis AX. Is.
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 increases or decreases a plurality of times according to the distance r from the optical axis AX 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 fluctuation width W1, and the difference between the maximum value and the minimum value of the second phase difference P2 is called the 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. The first region Z1 may be the central region of the intraocular lens including the reference point (eg, the optical axis AX, the center point) of the intraocular lens 10.
In the first embodiment, the first phase difference P1 and the second phase difference P2 are continuous and have a minimum value at the boundary BL between the first region Z1 and the first region Z2.

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 light beam when the first phase difference P1 and the second phase difference P2 are set to 0 does not have an image of an object 66 [cm] away from the eyeball 100. It is set to form an image on the macula 35 due to aberration. 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 the first embodiment, by changing the height of the injection surface 12 according to the distance r from the optical axis AX, the light rays passing through the first region Z1 and the light rays passing through the second region Z2 are positioned respectively. A phase difference (aberration) is added.
The first phase difference P1 added to the light beam passing through the intraocular lens 10 by the height H of the ejection surface 12 in the first region Z1 is, for example, an amount according to the Gaussian function with respect to the distance r from the optical axis. ..
Alternatively, the first phase difference P1 may be a quantity according to the sinc function with respect to the distance r from the optical axis or a function represented by the "power" of the sinc function.

Further, the first phase difference P1 may be a function based on the Gaussian function or the sin function described above, to which another function is added. That is, the first phase difference P1 may include a function based on 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 second phase difference P2 added to the light ray passing through the second region Z2 is expressed by using, for example, a distance r from the optical axis, an arbitrary constant a, an arbitrary constant b, and an arbitrary constant c. (1),
^{P2 = c × sin (ar 2} -b) ··· (1)
It is represented by.
Here, the constant a is a constant relating to the cycle of increase / decrease with respect to the distance r from the optical axis AX of the second phase difference P2, and when the unit of the distance r from the optical axis AX is [mm], 11.5 [ It is rad / mm ² ], and is preferably a value of 6 [rad / mm ² ] or more and 14 [rad / mm ² ] or less.

The constant b is a constant relating to the phase of increase / decrease with respect to the distance r of the second phase difference P2 from the optical axis AX. Since the sine function is a periodic function, the constant b has no upper and lower limits, but the value of the constant b is the first phase difference P1 and the second phase difference at the boundary BL between the first region Z1 and the second region Z2. Set so that P2 and P2 are continuous.
The constant c is, for example, 1.08 [rad].
The details of the value of the constant a and the value of the constant c will be described later.
The second phase difference P2 in the second region Z2 may be a function of the equation (1) plus another function. That is, the second phase difference P2 in the second region Z2 may include the function of the equation (1).

The maximum value of the first phase difference added to the light ray passing through the first region Z1 is the amount added to the light ray passing through the optical axis AX, which is +3.97 [rad]. On the other hand, the minimum value of the first phase difference is an amount in which the distance r from the optical axis AX is added to the light ray passing through the position of r1, and is -0.72 [rad]. Therefore, the fluctuation width W1 which is the difference between the maximum value and the minimum value of the first phase difference is 4.69 [rad].

The fluctuation width W2 of the second phase difference applied to the light rays passing through the second region Z2 is from the minimum value of -1.08 [rad] of the second phase difference P2 in the second region Z2 to the maximum value of 1.26 [rad]. Up to 2.34 [rad].
That is, in the first embodiment, the fluctuation width W1 of the first phase difference is set to be about twice as large as the fluctuation width W2 of the second phase difference.

The intraocular lens 10 of the first embodiment is the focal point of the optical system including the cornea 30, the anterior chamber 31, the intraocular lens 10, and the vitreous body 32 by adding the above phase differences P1 and P2 to the imaging light beam. The depth can be increased.
The increase in the depth of focus is partly due to the multifocal filter effect caused by the second phase difference P2 added to the transmitted light by the second region Z2 increasing or decreasing substantially periodically according to the distance r from the optical axis AX. by. That is, the intraocular lens 10 of the first embodiment divides a plurality of light rays (luminous flux) passing through the second region Z2 into a plurality of light fluxes. Then, the refractive powers (addition degree) of 0 diopter, +1.5 diopter, and +3 diopter are added to the plurality of luminous fluxes (-1st order diffracted light, 0th order diffracted light, and + 1st order diffracted light) divided into wave planes, respectively.

Further, in the intraocular lens 10 of the first embodiment, since the fluctuation width W1 of the first phase difference P1 is as large as about twice the fluctuation width W2 of the second phase difference P2, a plurality of lenses that have passed through the first region Z1. The depth of focus expansion effect due to the first phase difference P1 is added to the light beam (luminous flux).
As a result, in the intraocular lens 10 of the first embodiment, the multifocal effect by the plurality of light rays transmitted through the second region Z2 and the depth of focus expansion effect by the plurality of light rays transmitted through the first region Z1 are combined. Therefore, the depth of focus can be further expanded.

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. It is a graph which shows the simulation result in each diopter of the MTF (Modulation Transfer Function) at a predetermined spatial frequency (eg, about 50 [LP / mm]) of the image formed on the retina 33 in the vicinity.

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).

FIG. 3B shows three types of MTFs when the pupil diameter (diameter of the opening of the iris 36) is 3 mm, 4 mm, and 6 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 amount of defocus when the object is moved in the center line EX direction) is expressed in diopter [1 / m] units. For example, when an image of an object at infinity is formed at a position of 0 diopters on the horizontal axis, an image of an object 1 m away from the eyeball 100 is formed at a position of +1 dioptre on the horizontal axis.

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).
Comparing FIG. 3 (b) and FIG. 11 (a), the intraocular lens 10 of the present embodiment is 0, +1.5 in each pupil diameter as compared with the conventional single focus type intraocular lens. , And the MTF value can be increased at the focus position of +3 diopters. That is, it can be seen that the depth of focus of the eyeball 100 is expanded by the intraocular lens 10 of the present embodiment attached to the eyeball 100.
The MTF of the intraocular lens 10 of the present embodiment can obtain a good image for an object at infinity (0 diopter) and has a distance of about 66 cm and 33 cm with respect to the eyeball corresponding to +1.5 and +3 diopters. A good image can also be obtained for an object in.

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 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).
Even in the conventional double focus type intraocular lens, the MTF value can be increased especially at the defocus position of 0 and +2 diopters, but the MTF value at the defocus position of +1 diopter in the meantime is low.

On the other hand, in the intraocular lens 10 of the present embodiment, as shown in FIG. 3 (b), particularly when the pupil diameter is 3 mm, between each of the above-mentioned 0, +1.5 and +3 diopters. The MTF value can also be increased at the focus positions of +0.75 and +2.2 diopters. Further, the MTF value can be increased even at the focus position of +4 diopter. This is due to the effect of expanding the depth of focus due to the addition of the first phase difference P1 described above.
Since the pupil diameter of 3 mm is the pupil diameter in a bright place such as outdoors in the daytime, the intraocular lens 10 of the present embodiment is used in a bright place from a short distance of about 25 cm from the eyeball to infinity. A good image can be obtained for a wide range of objects.

In the above example, in the first region Z1, between the light ray passing near the optical axis AX and the light ray passing near the boundary BL with the second region Z2, as described above, there is a first phase difference. A phase difference (first phase difference P1) of 4.69 [rad], which is the difference between the maximum value and the minimum value, is added. However, this phase difference is not limited to the above-mentioned value of about 4.69 [rad], and if the phase difference is 2 [rad] or more or 3 [rad] or more, the above-mentioned depth of focus expansion effect can be obtained. Obtainable.

In the above example, the fluctuation width W1 of the first phase difference P1 is twice the fluctuation width W2 of the second phase difference P2, but the fluctuation width W1 of the first phase difference P1 is the second phase difference P2. If the fluctuation width is 1.5 times or more of W2, a large depth of focus expansion effect due to the first phase difference P1 can be obtained.
Further, in the present embodiment, if the fluctuation width W1 of the first phase difference P1 is 1.1 times or more the fluctuation width W2 of the second phase difference P2, a practically sufficient depth of focus expansion effect can be obtained. ..

When the pupil diameter is 3 mm, in order to obtain both the multiplex focus effect due to the second phase difference P2 and the depth of focus expansion effect due to the first phase difference P1, the first region Z1 is located inside the pupil diameter of 3 mm. And the second region Z2 need to exist in an appropriate area ratio.
For this purpose, the distance r1 from the optical axis AX at the boundary BL between the first region Z1 and the second region Z2 is preferably 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 cannot 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 it becomes impossible to obtain a sufficient multiplex effect.

The value of the constant a in the above equation (1) is a parameter that determines the amount corresponding to the period of the sin function in the equation (1), and therefore, the refractive power of the above-mentioned multifocal filter formed by the second region Z2. It becomes a parameter that determines.
The increase / decrease of the second phase difference P2 added by the second region Z2 is not limited to the shape by the sin function of the equation (1), and may be another shape.

However, in any case, the cycle of increase / decrease of the second phase difference P2 shall be 0.2 mm or more and 0.5 mm or less in the range where the distance r from the optical axis AX is from 1 mm to 1.5 mm. Is also good.
When the distance r is in the above range and the period of increase / decrease is in the above range, an appropriate refractive power (addition) can be added to the diffracted light generated from the second region Z2, and an appropriate multiplex focus can be applied. It can be effective.

Further, the cycle of increase / decrease of the second phase difference P2 may be 0.1 mm or more and 0.5 mm or less in the range of the distance from the optical axis AX from 1 mm to 2.5 mm.
When the distance r is in the above range and the period of increase / decrease is in the above range, an appropriate refractive power (addition) can be added to the diffracted light generated from the second region Z2, and more appropriate multiplexing can be performed. The focus effect can be exerted.

The value of the constant c in the above equation (1) is a parameter that determines the diffraction efficiency of the multifocal filter. When the increase / decrease of the second phase difference P2 follows the sin function of the equation (1), if the value of the constant c is 1.45 [rad], the wave plane is divided from the multifocal filter formed by the second region Z2. The light amounts of the three diffracted lights of light, 0th-order light, and + 1st-order light are approximately equal. Therefore, the value of the constant c is preferably a value of 0.75 to 1.75 [rad] in consideration of the balance of the amount of light of the -1st order light, the 0th order light, and the +1st order light.
When the constant c is in the above range, the fluctuation width W2 (maximum value-minimum value) of the second phase difference P2 is approximately 1.5 to 3.5 [rad].

As described above, the increase / decrease of the second phase difference P2 is not limited to the sin function of the equation (1), and may be based on another function. Even in that case, the fluctuation width W2 of the second phase difference P2 in the second region Z2 is 1.5 [rad] or more and 1.5 [rad] or more in the range where the distance r from the optical axis AX is 1 mm to 1.5 mm. It is preferably 3.5 [rad] or less.
When the distance r is in the above range and the fluctuation width W2 of the second phase difference P2 is in the above range, the amount of light of the -1st order diffracted light, the 0th order light diffracted light, and the +1st order diffracted light generated from the second region Z2. The balance of can be set to an appropriate value to effectively exert the multifocal effect.

Further, the fluctuation width W2 of the second phase difference P2 in the second region Z2 is 1.5 [rad] or more and 3.5 [rad] or more in the range where the distance r from the optical axis AX is 1 mm to 2.5 mm. ] The following is preferable.
When the distance r is in the above range and the fluctuation width W2 of the second phase difference P2 is in the above range, the amount of light of the -1st order diffracted light, the 0th order light diffracted light, and the +1st order diffracted light generated from the second region Z2. The balance of can be set to an appropriate value to make the multifocal effect more effective.

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.

(Modification example 1)
FIG. 4A is a diagram showing the relationship between the phase differences P1 and P2 added to the aberration-free imaging ray by the ejection surface 12 of the intraocular lens 10 of the modification 1 and the distance from the optical axis AX. is there. 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 intraocular lens 10 of the first modification, the change is gradual in the second region Z2 of the injection surface 12 where the second phase difference P2 increases as the distance r from the optical axis AX increases. The change is set to be steep at the portion where the second phase difference P2 decreases as the distance r increases. In other words, the second phase difference P2 has a change amount (differential amount with respect to the distance r) in the vicinity of one maximum value or a minimum value, particularly in the peripheral region 1.2 mm or more away from the optical axis, with respect to the optical axis side. It is asymmetrical on the side opposite to the optical axis side.
Due to the above-mentioned shape of the injection surface 12, in the intraocular lens 10 of the first modification, the amount of -1st-order diffracted light generated by dividing the wave surface from the second region Z2 can be made larger than the amount of +1-order diffracted light. ..

The second phase difference P2 added to the transmitted light by the second region Z2 is represented by the following equation (2) as an example.
^{P2 = c × sin (ar 2} -b)
+ D × sin (2ar ^2- b) ・ ・ ・ (2)
The first term on the right side of the equation (2) is the same as the right side of the equation (1). The second term on the right side of equation (2) is a term representing the so-called double frequency of the sin function of equation (1), and the value of the constant d is -0.1 to -0.4 times the value of the above constant c. It is a value of degree.
The second phase difference P2 may be obtained by further adding the 3 times frequency, 4 times frequency, and 5 times frequency of the above sine function to the equation (2).

FIG. 4B shows the vicinity of the 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 modification. It is a graph which shows the simulation result in each diopter of the MTF (Modulation Transfer Function) at a predetermined spatial frequency (eg, about 50 [LP / mm]) of the image formed on the retina 33 of the above. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

In the intraocular lens 10 of the first modification, as described above, the amount of -1st-order diffracted light generated by dividing the wave plane from the second region Z2 is set to be larger than the amount of +1st-order diffracted light. Therefore, the MTF at the focal position of 0 diopter (object at infinity) corresponding to the focal point of -1st order diffracted light is the focal point of +3 diopter (object 33 [cm] away from the eyeball 100) corresponding to the focal point of +1st order diffracted light. It can be higher than the MTF at the position. This effect is more pronounced when the pupil diameter is 4 mm or more, which includes more light rays passing through the second region Z2.
Similar to the intraocular lens 10 of the first embodiment shown in FIG. 3 (b), when the pupil diameter is 3 mm, the MTF value is increased even at the focus positions of +0.75 and +2.2 diopters. it can.

The second phase difference P2 of the intraocular lens 10 of the modification 1 does not necessarily have to follow the equation (2). However, since the second phase difference P2 added by the second region Z2 contains the component of the second term on the right side of the equation (2), the amount of light of the -1st order diffracted light and the +1st order diffracted light generated from the second region Z2. The ratio of can be changed. As a result, it is possible to realize an intraocular lens 10 that emphasizes a distant view and an intraocular lens 10 that emphasizes a closer view.

(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 from the optical axis AX. is there. 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.

In the intraocular lens 10 of the second modification, in the second region Z2 of the injection surface 12, the distance r from the optical axis AX is larger than the fluctuation width W21 of the second phase difference P2 in the range of 1 to 1.5 mm. The fluctuation width W22 of the second phase difference P2 in the range where r is 2 to 2.5 mm is larger.
That is, the intraocular lens 10 of the modified example 2 exhibits a larger multiplex effect than the intraocular lens 10 of the modified example 1 when the pupil diameter exceeds 4 mm.

FIG. 5B shows each diopter 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 in. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

In the intraocular lens 10 of the modified example 2, compared to the intraocular lens 10 of the modified example 1 shown in FIG. 4 (b), when the pupil diameter is 6 mm (when it exceeds 4 mm), it becomes 0 diopter and +3 diopter. The MTF at the corresponding focus position is increasing. That is, when the pupil diameter exceeds 4 mm, a larger multiplex effect is exhibited.
This is suitable for, for example, the intraocular lens 10 when it is necessary to emphasize the distant view and the near view in an environment where the pupil diameter exceeds 4 mm, that is, in a slightly dark environment such as indoors or at dusk.

(Modification 3)
FIG. 6A 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 from the optical axis AX. is there. Since the configuration of the intraocular lens 10 of the modification 3 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.

In the intraocular lens 10 of the third modification, in the second region Z2 of the injection surface 12, the distance r from the optical axis AX is larger than the fluctuation width W22 of the second phase difference P2 in the range of 2 to 2.5 mm. The fluctuation width W21 of the second phase difference P2 in the range where r is 1 to 1.5 mm is larger.
That is, the intraocular lens 10 of the modified example 2 exhibits a larger multiplex effect than the intraocular lens 10 of the modified example 1 when the pupil diameter is 4 mm or less.

FIG. 6B shows each diopter 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 in. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

In the intraocular lens 10 of the modified example 3, compared to the intraocular lens 10 of the modified example 1 shown in FIG. 4 (b), when the pupil diameter is 4 mm or less, the focus position corresponds to 0 diopter and +3 diopter. MTF is increasing. That is, when the pupil diameter is 4 mm or less, a larger multiplex effect is exhibited.

(Modification example 4)
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 modified example 4 and the distance from the optical axis AX. is there. Since the configuration of the intraocular lens 10 of the modified example 4 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 intraocular lens 10 of the modified example 4, the first phase difference P1 decreases monotonically as the distance r from the optical axis AX increases, and the phase differences P1 and P2 have a distance r from the optical axis AX r2. The first minimum value MP2 is taken at the position where (about 0.4 mm) is obtained. In the region where the distance r from the optical axis AX is r2 or more, the second phase difference P2 repeatedly increases and decreases according to the distance r from the optical axis AX, but in the region where the distance r from the optical axis AX is 2 mm or more, The fluctuation width W22 of the second phase difference P2 gradually decreases as the distance r increases.
The fluctuation width W21 of the second phase difference P2 in the range where the distance r from the optical axis AX is 1 to 1.5 mm is about 2.2 [rad].

In the intraocular lens 10 of the modified example 4, the cycle of increase / decrease with respect to the distance r of the second phase difference P2 from the optical axis AX is shorter than that of the first embodiment and other modified examples. Therefore, the refractive powers (addition degree) of −3.5 diopters and +3.5 diopters are added to the -1st order diffracted light and the + 1st order diffracted light generated from the second region Z2, respectively.

In the intraocular lens 10 of the modified example 4, the shapes and arrangements of the incident surface 11 and the reference surface 13 are such that when the phase differences P1 and P2 are 0, the eyeball 100 loaded with the intraocular lens 10 is the eyeball 100. It is set to form an image of an object 30 [cm] away from the macula 35 without aberration. Therefore, for the -1st order diffracted light to which the addition degree of −3.5 dioptres is added in the second region, an image of an object at almost infinity is formed on the macula 35 in the eyeball 100.

Further, in the intraocular lens 10 of the modified example 4, in the range where the distance r from the optical axis AX in the second region Z2 is in the range of 0.4 to 1 mm, the second position is increased according to the increase in the distance r from the optical axis AX. The change is steep in the portion where the phase difference P2 increases, and the change is set to be gradual in the portion where the second phase difference P2 decreases as the distance r increases. Such a change in the second phase difference P2 can be obtained, for example, by setting the value of the constant d to a value of about +0.1 to +0.4 times the value of the constant c in the above equation (2). ..

As a result, in the intraocular lens 10 of the modified example 4, the pupil diameter is particularly small (when the pupil diameter is 4 mm or less, the amount of +1st order light generated from the second region Z2 is larger than the amount of -1st order light. Therefore, it is possible to realize the intraocular lens 10 having excellent resolution of a near view, particularly in a relatively bright environment.

FIG. 7B shows each diopter 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 modified example 4. It is a graph which shows the simulation result in. The conditions of the simulation are the same as those shown in FIG. 3 (b) above, but FIG. 7 (b) also shows the MTF when the pupil diameter is 2 mm.

In the intraocular lens 10 of the modified example 4, compared to the intraocular lens 10 of the modified example 1 shown in FIG. 7 (b), when the pupil diameter is 4 mm or less, the focus position corresponds to 0 diopter and +3 diopter. MTF is increasing. That is, when the pupil diameter is 4 mm or less, a larger multiplex effect is exhibited.

As in the above-mentioned modifications 2 to 4, the fluctuation widths W21 and W22 of the second phase difference P2 added to the light rays passing through the second region are increased or decreased according to the distance from the optical axis AX. As a result, the multifocal effect due to the second phase difference P2 can be changed according to the pupil diameter. This makes it possible to realize an intraocular lens 10 in which the object distance at which a good image can be obtained changes according to the brightness of the surroundings.

(Modification 5)
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 5 and the distance from the optical axis AX. is there. Since the configuration of the intraocular lens 10 of the modified example 5 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 intraocular lens 10 of the modified example 5, in the second region Z2 of the injection surface 12, the fluctuation width W21 of the second phase difference P2 in the range where the distance r from the optical axis AX is about 0.5 mm to 1.5 mm is , 2.2 [rad]. However, the fluctuation width W22 of the second phase difference P2 in the range where the distance r from the optical axis AX is 2 to 2.5 mm is 0.5 [rad] or less.
Therefore, in the range where the distance r from the optical axis AX in the second region Z2 is 2 to 2.5 mm, the multiplex effect is hardly exhibited.

FIG. 8B shows each diopter 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 modified example 5. It is a graph which shows the simulation result in. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

In the intraocular lens 10 of the modified example 5, as described above, the multifocal effect is weakened at a pupil diameter of about 4 mm or more, so that the MTF at 0 diopter and +3 diopter becomes a low value. However, the MTF at +1.5 diopters, which corresponds to the focal position of the 0th order diffracted light from the second region Z2, can be increased.
That is, the intraocular lens 10 of the modified example 5, like the intraocular lens 10 of the first embodiment and other modified examples, realizes an expansion of the depth of focus, and the eyeball 100 is particularly in a dark environment. A near view at a predetermined distance (for example, about 66 [cm]) can be satisfactorily imaged on the macula 35.

In the above first embodiment and the intraocular lens 10 of each modification, the phase differences P1 and P2 are determined according to the distance r from the optical axis AX, that is, they have a shape rotationally symmetric with respect to the optical axis AX. There is. 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 second phase difference 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 (3). The second phase difference P2 may be determined by replacing it with r'.
r'= r × {1 + g · sin (2θ－φ)} ・ ・ ・ (3)
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.
Similarly, when the first phase difference P1 added by the first region Z1 is determined based on a function such as a Gaussian function or a sinc function with respect to the distance r from the optical axis AX as described above, those functions are also used. As the argument of, the above r'may be used instead of the distance r.

(Modification 6)
FIG. 9A is a top view of the intraocular lens 10a of the modified example 6.
In the above-described first embodiment and each modification, the first region Z1 is a region centered on the optical axis AX of the intraocular lens 10a, but in the intraocular lens 10a of the modification 6, the boundary BL The center of the first region Z1 surrounded by is eccentric from the optical axis AX. Therefore, each point on the boundary BL is not equidistant from the optical axis AX.

(Modification 7)
FIG. 9B is a top view of the intraocular lens 10b of the modified example 7. Also in the intraocular lens 10b of the modified example 7, the center of the first region Z1 surrounded by the boundary BL is located at a position eccentric from the optical axis AX, and the optical axis AX intersects with the first region Z1. Not.
The intraocular lenses 10 of the modified examples 6 and 7 are suitable for correcting astigmatism of intensity caused by, for example, the shape of the cornea 30.

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 first 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 first phase difference P1 which increases monotonically according to the distance r from the optical axis takes the first maximum value. Become.

(Modification 8)
FIG. 10A shows the phase differences P1 and P2 and the optical axis AX in the intraocular lens 10 of the modified example 8 having a shape in which the first phase difference P1 monotonically increases with the distance r from the optical axis. It is a figure which shows the relationship with the distance from. Since the configuration of the intraocular lens 10 of the modified example 8 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 intraocular lens 10 of the modified example 8, the height H of the ejection surface 12 is lower than the reference surface 13 in the vicinity of the optical axis AX, and the negative phase difference P1 (with respect to the light rays transmitted in the vicinity of the optical axis AX). As an example, add about -25 [rad]).
As the distance r from the optical axis AX increases, the first phase difference P1 added to the transmitted light by the first region Z1 increases monotonically. Then, at the position (boundary BL) where the distance r is 1.4 mm (distance r1), the phase differences P1 and P2 take the maximum value MM1.

In the second region Z2 farther from the optical axis AX than the boundary BL, the second phase difference P2 decreases as the distance r increases, and at the position where the distance r is 1.6 mm, the minimum value MP2 adjacent to the maximum value MM1 Take. The difference W0 between the maximum value MM1 and the minimum value MP2 is, for example, about 1.4 [rad]. That is, the second phase difference P2 changes by about 1.4 [rad] from the maximum value MM1 to the minimum value MP2. Hereinafter, this value will be referred to as the first value W0.
The second phase difference P2 repeats increasing and decreasing even in a region where the distance r from the optical axis AX is larger than the position of the minimum value MP2 described above. As described above, the fluctuation range is the fluctuation range W2.
In the intraocular lens 10 of the modified example 8, the fluctuation width W2 is twice or more the first value W0. Further, similarly to the above-described first embodiment and each modification, the fluctuation width W1 is 2 [rad] or more, or 3 [rad] or more larger than the fluctuation width W2.

FIG. 10B shows each diopter 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 modified example 8. It is a graph which shows the simulation result in. The conditions of the simulation are the same as those of the simulation shown in FIG. 3 (b) above.

In the intraocular lens 10 of the modified example 8, compared to the conventional single focus type intraocular lens whose MTF simulation result is shown in FIG. 11 (b), at each pupil diameter, particularly at the focus position of +1 to +3 diopters. The value of MTF can be increased. That is, it can be seen that the depth of focus of the eyeball 100 is expanded by the intraocular lens 10 of the present embodiment attached to the eyeball 100.

Further, as compared with the simulation result of MTF using the conventional double focus type intraocular lens shown in FIG. 11B, the intraocular lens 10 of the modified example 8 has MTF at continuous focus positions instead of discrete ones. It can be seen that the value of can be increased.
The signs of the phase differences P1 and P2 of the intraocular lens 10 of the modified example 8 shown in FIG. 11A may also be inverted (inverted in the vertical axis direction).

(Supplementary explanation regarding the first embodiment and each modification)
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 continuously change according to the distance r from the optical axis. However, the first 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 first phase difference P1 and the first phase difference P2 are not only continuous but also smooth at the boundary BL between the first region Z1 and the second region Z2. It may be continuous with. In this case, it is possible to realize an intraocular lens 10 that provides a clearer field of view by reducing scattered light generated by abrupt changes in phase differences P1 and P2 near the boundary BL.

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

Alternatively, by providing a refractive index fluctuation portion in the first region Z1 and the second region Z2 of the intraocular lens 10 in which the refractive index changes with respect to the distance r from the optical axis AX, the above-mentioned phase difference P1 and P2 can be added. 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 difference between the light rays passing through the first region Z1 and the second region Z2 has the above-mentioned 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 100, and is worn on the eyeball 100. It includes a first region Z1 and a second region Z2 that transmit light to form an image on the retina. Then, the first region Z1 adds a first phase difference P1 that monotonically increases or decreases monotonically according to the distance from the optical axis AX of the ophthalmic lens 10 to the light rays passing through the first region Z1, and the second region Z1. Z2 adds a second phase difference P2 that increases or decreases a plurality of times according to the distance from the optical axis AX to the light rays passing through the second region Z2. The fluctuation width W1 of the first phase difference P1 is 1.1 times or more the fluctuation width W2 of the second phase difference P2.
With this configuration, the multifocal effect of a plurality of light rays transmitted through the second region Z2 and the depth of focus expansion effect of a plurality of light rays transmitted through the first region Z1 are combined, and the depth of focus is expanded for ophthalmology. The lens 10 can be realized.

(2) The first region Z1 is arranged so as to intersect the optical axis AX, and 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. Moreover, the minimum value MP1 or the maximum value may be taken at the boundary BL, and the distance from the optical axis AX to the boundary BL may be 0.4 mm or more and 0.8 mm or less.
With this configuration, when the pupil diameter is 3 mm in a bright environment, the multiple focus effect due to the second phase difference P2 and the depth of focus expansion effect due to the first phase difference P1 are exhibited in a well-balanced manner, and the depth of focus is further expanded. The ophthalmic lens 10 can be realized.

(3) The cycle of increase / decrease of the second phase difference P2 may be 0.2 mm or more and 0.5 mm or less in the range of the distance r from the optical axis AX from 1 mm to 1.5 mm.
Further, the cycle of increase / decrease of the second phase difference P2 may be 0.1 mm or more and 0.5 mm or less in the range of the distance r from the optical axis AX from 1 mm to 2.5 mm.
When the distance r is in the above range and the period of increase / decrease is in the above range, an appropriate refractive power (addition) can be added to the diffracted light generated from the second region Z2, and an appropriate multiplex focus can be applied. It can be effective.
(4) The fluctuation width W1 of the first phase difference P1 may be 1.5 times or more the fluctuation width W2 of the second phase difference P2. With this configuration, the effect of expanding the depth of focus by the plurality of light rays transmitted through the first region Z1 can be further increased.

(5) The fluctuation width W2 of the second phase difference P2 may be 1.5 [rad] or more and 3.5 [rad] or less in the range of the distance r from the optical axis AX from 1 mm to 1.5 mm. .. With this configuration, it is possible to more effectively exert the multiple focus effect by the plurality of light rays transmitted through the second region Z2, particularly in a bright environment where the pupil diameter is about 3 mm.
(6) The fluctuation width W2 of the second phase difference P2 may be 0.5 [rad] or less in the range where the distance r from the optical axis AX is 2 mm or more. With this configuration, while realizing the expansion of the depth of focus, an image of an object at a predetermined distance from the eyeball 100 is satisfactorily imaged on the macula 35, especially in a dark environment where the pupil diameter is about 5 mm or more. be able to.

(7) The fluctuation width W2 of the second phase difference P2 may be 1.5 [rad] or more and 3.5 [rad] or less in the range of the distance r from the optical axis AX from 1 mm to 2.5 mm. .. With this configuration, the balance of the amount of light of the -1st order diffracted light, the 0th order diffracted light, and the + 1st order diffracted light generated from the second region Z2 can be set to an appropriate value, and the multiplex effect can be more effectively exhibited. it can.
(8) The first phase difference P1 may be configured to change continuously in the first region Z1. As a result, it is possible to realize an ophthalmic lens that reduces scattered light generated by a sudden change in phase difference and provides a clearer field of view.

(9) The first region Z1 is formed between a light ray passing near the optical axis AX in the first region Z1 and a light ray passing near the boundary BL with the second region Z2 in the first region Z1. It may be configured to add a phase difference of [rad] or more. With this configuration, the effect of expanding the depth of focus by the plurality of light rays transmitted through the first region Z1 can be further increased.

(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 such as the above phase differences P1 and P2 (eg, design conditions) into a lens shape.

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

Claims (25)

1. An ophthalmic lens that is 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 adds a first phase difference that monotonically increases or decreases monotonically depending on the distance from the optical axis of the ophthalmic lens to the light rays passing through the first region.
   The second region adds a second phase difference that increases or decreases a plurality of times according to the distance from the optical axis to the light rays passing through the second region.
   An ophthalmic lens in which the fluctuation range of the first phase difference is 1.1 times or more the fluctuation range of the second phase difference.
2. In the ophthalmic lens according to claim 1.
   The first region intersects the optical axis and
   The first phase difference and the second phase difference are continuous at the boundary between the first region and the second region, and take a minimum value or a maximum value at the boundary.
   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.
3. In the ophthalmic lens according to claim 1 or 2.
   An ophthalmic lens in which the cycle of the increase / decrease of the second phase difference is 0.2 mm or more and 0.5 mm or less in a range of a distance from the optical axis of 1 mm to 1.5 mm.
4. In the ophthalmic lens according to any one of claims 1 to 3.
   An ophthalmic lens in which the cycle of the increase / decrease of the second phase difference is 0.1 mm or more and 0.5 mm or less in a range of a distance from the optical axis of 1 mm to 2.5 mm.
5. In the ophthalmic lens according to any one of claims 1 to 4.
   An ophthalmic lens in which the fluctuation width of the first phase difference is 1.5 times or more the fluctuation width of the second phase difference.
6. In the ophthalmic lens according to any one of claims 1 to 5.
   An ophthalmic lens having a fluctuation range of the second phase difference of 1.5 [rad] or more and 3.5 [rad] or less in a range of 1 mm to 1.5 mm from the optical axis.
7. In the ophthalmic lens according to any one of claims 1 to 6.
   An ophthalmic lens having a fluctuation range of the second phase difference of 0.5 [rad] or less in a range of 2 mm or more from the optical axis.
8. In the ophthalmic lens according to any one of claims 1 to 6.
   An ophthalmic lens having a fluctuation range of the second phase difference of 1.5 [rad] or more and 3.5 [rad] or less in a range of 1 mm to 2.5 mm from the optical axis.
9. In the ophthalmic lens according to claim 1.
   The first region intersects the optical axis and adds the first phase difference that monotonically increases or decreases depending on the distance from the optical axis to the light rays passing through the first region.
   The first phase difference and the second phase difference are continuous at the boundary between the first region and the second region, and take a minimum value or a maximum value at the boundary.
   An ophthalmic lens having a distance from the optical axis to the boundary of 0.4 mm or more and less than 1.5 mm.
10. In the ophthalmic lens according to claim 9.
    The second phase difference is
    The first value changes between the boundary that is the minimum value or the maximum value and the maximum value or the minimum value that is adjacent to the side away from the optical axis with respect to the boundary.
    An ophthalmic lens in which the fluctuation range of the second phase difference is at least twice the first value.
11. The ophthalmic lens according to any one of claims 1 to 10.
    An ophthalmic lens in which the first phase difference changes continuously within the first region.
12. In the ophthalmic lens according to any one of claims 1 to 11.
    The first region is 3 [rad] or more between a light ray passing near the optical axis of the first region and a light ray passing near the boundary with the second region of the first region. An ophthalmic lens that adds the phase difference of.
13. The ophthalmic lens according to any one of claims 1 to 12.
    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.
14. The ophthalmic lens according to any one of claims 1 to 12.
    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.
15. The ophthalmic lens according to any one of claims 1 to 14.
    An ophthalmic lens in which the fluctuation range of the second phase difference increases or decreases according to the distance from the optical axis.
16. The ophthalmic lens according to any one of claims 1 to 15.
    The second phase difference is a distance r from the optical axis, arbitrary constant a, using any of the constants b, containing a component represented by sin (ar 2 ^-b), ophthalmic lenses.
17. In the ophthalmic lens according to claim 14.
    It said second phase difference, further, sin comprising a component represented by ^(2ar 2-2b), ophthalmic lenses.
18. The ophthalmic lens according to any one of claims 1 to 17.
    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.
19. The ophthalmic lens according to any one of claims 1 to 18.
    An ophthalmic lens in which the first phase difference and the second phase difference are rotationally symmetric with respect to the optical axis.
20. The ophthalmic lens according to any one of claims 1 to 18.
    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.
21. The ophthalmic lens according to any one of claims 1 to 20.
    The ophthalmic lens is an intraocular lens that is loaded into the eyeball instead of the crystalline lens.
22. The ophthalmic lens according to any one of claims 1 to 20.
    The ophthalmic lens is an ophthalmic lens which is an implantable contact lens loaded between the iris and the crystalline lens.
23. The ophthalmic lens according to any one of claims 1 to 20.
    The ophthalmic lens is an ophthalmic lens which is a corneal inlay or a corneal inlay loaded into the cornea.
24. The ophthalmic lens according to any one of claims 1 to 20.
    The ophthalmic lens is an ophthalmic lens which is a contact lens that comes into contact with the cornea.
25. 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 24 is manufactured by a processing apparatus.
    

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