**Purpose**:
Choroid geometry and swelling have been proposed to contribute to ocular pathologies. Thus, it is important to understand how the choroid may impact the optic nerve head (ONH) biomechanical environment. We developed a finite element model to study how acute choroidal swelling and choroid geometry affect ONH deformation.

**Methods**:
We developed two geometric models of the ONH: one with a “blunt” choroidal insertion and another with a “tapered” choroid insertion. We examined how choroidal volume changes (2.1–14.2 μL, estimated to occur during the ocular pulse) impact biomechanical strain in three tissue regions: the prelaminar neural tissue, lamina cribrosa, and retrolaminar neural tissue. Then, we performed a sensitivity analysis to understand how variation in ONH pressures, tissue material properties, and choroidal swelling influenced the peak tissue strains.

**Results**:
Choroidal swelling in the blunt choroid geometry had a large impact on the strains in the prelaminar neural tissue, with magnitudes comparable to those expected to occur due to an IOP of 30 mm Hg. Choroidal swelling in the tapered choroid geometry also affected strains but to a lesser extent compared to the blunt geometry. A sensitivity analysis confirmed that choroidal swelling was more influential on prelaminar neural tissue strains in the blunt choroid geometry.

**Conclusions**:
Choroid anatomy and swelling can interact to play an important role in prelaminar neural tissue deformation. These findings suggest that the choroid may play an important, and previously unappreciated, role in ONH biomechanics, and motivate additional research to better define the in vivo effects of choroidal volume change.

^{1,2}Deviations in these loads from the physiological range are proposed to play a role in several ocular pathologies, including glaucoma, idiopathic intracranial hypertension (IIH), and space flight–associated neuro-ocular syndrome (SANS).

^{3,4}For example, elevated IOP is the primary risk factor for glaucoma, the second leading cause of blindness worldwide.

^{5}Similarly, changes in ICP play a role in visual impairment in IIH. The alteration in ICP magnitude, or its daily variation, is also thought to contribute to visual impairment in SANS. These loads cause deformations of ONH tissues which, if they exceed some unknown threshold, may lead to either direct (e.g., mechanical insult) or indirect (e.g., occlusion of blood flow or astrocyte activation) injury and damage to retinal ganglion cells (RGCs). The lamina cribrosa is of special interest in this context because it undergoes remodeling and is an early and major site of RGC injury in glaucoma.

^{6–8}However, to date, few FE models have considered the effects of the choroid on ONH biomechanics,

^{9,10}and in general, the impact of changes in choroidal volume and geometry on ONH deformation is largely unknown.

^{11}Specifically, Nagia et al.

^{11}showed that choroidal thickness is larger in NAION patients compared to controls. We thus wondered how choroidal volume changes would affect ONH tissue biomechanics. We hypothesized that choroidal swelling would lead to appreciable deformations of the ONH, which in turn could lead to crowding of structures at the optic disc or activation of remodeling pathways leading to RGC injury. Such biomechanical changes could induce a compartment syndrome in the prelaminar neural tissue or potentially cause nonphysiological loading on ONH tissues.

^{7,12,13}We also incorporated a single central retinal vessel to represent both the central retinal artery and vein. This idealized vessel allowed us to simulate the effects of blood pressure through specification of a mean arterial pressure (MAP). Finally, these ocular models included two novel components: the choroidal tissue and Bruch's membrane (Fig. 1).

^{14,15}specifically an average thickness of 136 μm over the first 500 μm adjacent to the scleral canal margin,

^{15}increasing anteriorly to 250 μm at 2000 μm away from the scleral canal margin.

^{14–16}After this maximum thickness was achieved, the choroid thinned anteriorly until it reached the ora. Similarly, farther than 2000 μm away from the scleral canal margin, the thickness of the PLNT decreased anteriorly. The rate of choroidal thinning with anterior location was based upon a constraint on the total volume of the choroid. Specifically, the total choroidal volume was calculated based on ocular parameters such as eye diameter and maximum choroidal thickness

^{17}to obtain a choroidal blood volume of 192 μL. It has been suggested that the choroidal blood volume represents up to 75% of the total choroidal volume; thus, the unloaded choroidal volume, that is, the combination of the blood and tissues, was taken as 256 μL.

^{17}Lastly, our model included an idealized Bruch's membrane between the choroid and PLNT, spanning from the termination of the choroid at the ONH to the ora. We modeled this structure as having a uniform thickness of 3 μm.

^{11}and Rhodes et al.

^{16}The choroidal thickness was taken from optical coherence tomography (OCT) measurements as 61, 102, and 136 μm at locations 125, 375, and 750 μm away from the scleral canal opening, respectively. Similar to the blunt geometry model, we ensured that the maximum choroidal thickness was 250 μm at a distance 2000 μm from the scleral canal margin. The unloaded choroidal volume was again set to be 256 μL; that is, the baseline choroidal volume was identical between the two geometries, and a uniform Bruch's membrane was included in the model between the choroid and PLNT. In our study, the major geometric change between the blunt and tapered models was the choroid. These changes in the choroidal geometry also lead to secondary, minor changes in the length of the Bruch's membrane and the surrounding PLNT.

^{18}while our FE simulations were performed in the FE solver FEBio.

^{19}The geometry of the eye for FE analysis was treated as axisymmetric—represented as a 3° wedge about an axis of symmetry passing through the central retinal vessel. This wedge geometry represents an axisymmetric model in the FEBio solver. In brief, we performed a convergence study, considering our outcome measures (peak first and third principal strains) at each loading condition within the lamina cribrosa, optic nerve, and PLNT. Our production meshes resulted in less than a 5% relative error in the peak first and third principal strain in each tissue region compared to the most refined mesh. The production meshes contained 106,671 elements for the blunt geometry and 101,368 elements for the tapered geometry.

^{20,21}The fiber orientation, alignment, and material properties were based on an earlier published FE model. In brief, the preferred fiber orientation (

*θ*) and degree of alignment (

_{P}*k*) were defined for each tissue region. In the posterior sclera, the fibers were planar isotropic (

_{f}*k*= 0) lying within the local tangent plane of the sclera.

_{f}^{22,23}In the peripapillary sclera and annular ring, the fibers were oriented circumferentially around the scleral canal with fiber concentration factors (

*k*) of 0.85 and 1.85, respectively.

_{f}^{23}Similar to the posterior sclera, the pia mater and dura mater fiber distributions were taken as planar isotropic (

*k*= 0) with the fibers lying within the local tangent planes.

_{f}*c*), representing all the constituents except the collagen fibers (e.g., proteoglycans, cells, elastin). We then defined two coefficients that described the stiffness of the collagen fibers (

_{1}*c*and

_{3}*c*). Lastly, we defined a bulk modulus to enforce tissue incompressibility (

_{4}*K*= 100 MPa).

*E*) and the Poisson ratio (

*ν*). Young's modulus values were adopted from previous experimental and FE studies of the posterior eye.

^{6,24}Consistent with earlier FE models, we assumed that the neural tissue (e.g., the PLNT and optic nerve) were partially compressible (

*ν*= 0.45) while the remaining tissues were nearly incompressible (

*ν*= 0.49).

^{6,13}

^{25–27}(

*ΔV*, so that no choroidal swelling is represented by

*ΔV*= 0 μL. This condition was considered as our baseline.

^{28}who reported ocular pulse amplitudes of 0.9 mm Hg (minimum), 3 mm Hg (median), and 7.2 mm Hg (maximum). Using the empirical correlation of Silver and Geyer,

^{29,30}the above IOP changes correspond to ocular volume changes of 2.1, 6.5, and 14.2 μL, respectively. Therefore, associated with the ocular pulse we assumed the choroid expands between 2.1 and 14.2 μL in a healthy individual. For comparison, we also simulated the impact of elevated IOP (IOP = 30 mm Hg) without any choroidal swelling (

*ΔV*= 0 μL). For these simulations, ICP and MAP remained constant. In summary, this resulted in five separate loading conditions: (1) IOP = 15 mm Hg and 0 μL of choroidal swelling, (2) IOP = 15 mm Hg with 2.1 μL of choroidal swelling, (3) IOP = 15 mm Hg with 6.5 μL of choroidal swelling, (4) IOP = 15 mm Hg with 14.2 μL of choroidal swelling, and (5) IOP = 30 mm Hg with 0 μL of choroidal swelling. For simplicity, we refer to these loading conditions by the primary parameter being investigated: (1) IOP = 15 mm Hg, (2)

*ΔV*= 2.1 μL, (3)

*ΔV*= 6.5 μL, (4)

*ΔV*= 14.2 μL, and (5) IOP = 30 mm Hg.

^{31}The peak first and third principal strains were specifically defined as the 95th and 5th percentile strain within each region of interest, respectively. This definition excludes numerical outlier values of strain, due, for example, to poorly formed FE mesh elements. We focused our analysis on three regions: the PLNT within 1 mm of the lamina cribrosa; the lamina cribrosa; and the retrolaminar neural tissue (RLNT). The RLNT was defined as the optic nerve region within 1 mm of the posterior lamina cribrosa surface.

*α*= 0.05), and a two-sample Kolmogorov-Smirnov test to perform a pairwise comparison between each condition with Bonferroni correction for multiple comparisons (

*α*= 0.05 / 5 = 0.01).

^{32}LHS allowed us to define the specific distribution and range of each input parameter. Here, all tissue material properties and choroidal swelling were assumed to follow uniform distributions, as described in Table 1. The IOP, ICP, and MAP followed normal distributions based on ranges reported for healthy individuals in the upright position.

^{13,33,34}

^{13}This resulted in a total of 400 unique simulations per geometry in our sensitivity analysis. The outcome measures we considered were the peak first and third principal strains in the PLNT, lamina cribrosa, and RLNT. The input parameters were related to each outcome measure by a partial rank correlation coefficient ranging between −1 and 1. The closer to 1 a value is in either the positive or negative direction, the greater the linear relationship between that input parameter and the resulting peak first and third principal strains. A value of zero means there is no linear relationship between the input parameter and outcome measure. After the partial rank correlation coefficients were calculated, the input parameters were ranked based on their proximity to ±1, with the strongest correlation given a rank of 25 and the lowest correlation given a rank of 1. These ranked scores were calculated for each tissue region (e.g., PLNT, lamina cribrosa, and RLNT) and outcome measure (i.e., peak first and third principal strains), added together, and then normalized to the maximum possible score (150, i.e., maximum score of 25 in three tissue regions with two strain measures). An input parameter with the most influence across each tissue region would thus have a normalized score closer to 1 while the least influential parameter would have a normalized score near 0.

*ΔV*= 0 μL) we found no significant differences in the first and third principal strain distributions between the blunt and tapered choroid geometries in any tissue region (Table 2). However, we did find a significant difference in the third principal strain distributions in the PLNT for a choroidal swelling of 2.1 μL, but no significant differences in strain distributions in the lamina cribrosa or RLNT. For more choroidal swelling (

*ΔV*= 14.2 μL), there were significant changes in the first and third principal strain distributions in the PLNT between the blunt and tapered choroid geometries (Table 2). Finally, we noted that the strain distributions in the RLNT were not significantly different when IOP was elevated to 30 mm Hg (

*ΔV*= 0 μL) when comparing the blunt and tapered choroid geometries. However, elevating IOP did significantly change the third principal strain distribution within the PLNT and the lamina cribrosa (Table 2), indicating that choroid geometry interacts with elevated IOP to affect ONH tissue strains.

^{14}presence of a specific disease (i.e., NAION, glaucoma),

^{11,15}or ethnicity.

^{16}These future modeling studies may provide additional insights into the choroid's role in influencing ONH biomechanics.

^{7,8,12}It is commonly accepted that patients with an IOP of 30 mm Hg would be monitored closely due to an increased risk of developing a glaucomatous optic neuropathy; and thus significant, chronic choroidal swelling a priori could be of potential clinical importance depending on the individual's choroidal anatomy.

^{35,36}Since the ocular pulse is physiological, it is unlikely that such fast, transient stimuli would induce pathology without additional factors. However, persistent choroidal swelling above a certain physiological range could activate ONH mechanosensitive cells and result in pathology. From glaucoma studies, we hypothesize that a sustained deformation due to elevation of IOP can stimulate mechanically sensitive cells in the ONH, thereby initiating a biological cascade that results in permanent pathological remodeling.

^{37–39}The impact of persistent choroidal swelling on the surrounding tissues in the ONH thus merits further investigation, particularly since the choroid is a dynamic tissue that is responsive to multiple stimuli, including pharmacological and positional challenges among others.

^{40–44}

^{13}giving us confidence in the model. Our results are also generally consistent with previous work showing that other ocular anatomic features (e.g., scleral thickness) influence OHN strains due to elevated IOP.

^{6,12}Interestingly, clinical evidence suggests that IOP and choroidal volume themselves may be linked.

^{45,46}A reduction of IOP via trabeculectomy or deep sclerectomy has been shown to cause a subsequent change in choroidal thickness. Thus, studies aimed at understanding the interaction of IOP and choroid volume and their combined influence on ONH biomechanics will be critical in the future.

^{11}These geometries used average values across several studies, yet represented only two potential geometries. Nonetheless, we believe these data provide us valuable and novel insight into how variations in choroid geometries and swelling influence ONH deformation and motivate further study of this effect. In addition, we used an axisymmetric geometric model to represent the complex anatomy of the eye, which by definition cannot include variations in choroidal thickness between the inferior, temporal, superior, and nasal regions.

^{11,14,15}Thus, we used an average value to represent choroidal thickness at various locations and aimed to match an overall calculated volume of the choroid for each geometry. To capture more realistic aspects of choroidal anatomy, it will be necessary in future to develop full three-dimensional, data-driven geometric models of the posterior eye with varying choroidal thicknesses. Thus, while these simplifications are a limitation of the present study, they also clearly highlight the need to collect additional individual-specific experimental data about choroid anatomy and swelling throughout the entire eye.

^{47,48}Since such effects were beyond the scope of the present study and research question, we have not assessed how variations in the location, degree, and width of circumferentially aligned collagen fibers impacts deformation at the ONH. Understanding how this variation affects deformation could provide additional insight into ONH biomechanics.

^{49}estimated an average volume change of 3.98 μL (min 1.87 μL and max 7.19 μL) over a cardiac cycle, which encompasses the two smaller volume changes we simulated in the present study (2.1 μL and 6.5 μL). A drawback of these estimates is they assumed a simplified choroidal shape and estimated blood flow into the choroid. We also estimated choroidal volume changes of 13.1 μL from reported OCT measurements,

^{50}which is similar to the largest choroidal volume change we specified in our models (14.2 μL). An important limitation of using OCT measurements is the limited field of view of the choroid, so that OCT measurements must be extrapolated to the entire choroid. However, these independent studies led to choroidal volume change estimates comparable to the values we derived from ocular pulse amplitudes, increasing confidence in our approach. Future studies should seek to use more directly measured changes in choroidal volume under various conditions, which we hope will arise from advances in OCT and better understanding of ocular blood flow. In addition, our simulations do not directly account for the changing IOP over the ocular pulse; that is, we applied a constant IOP within each simulation. This is a limitation of the quasi-steady modeling framework used in the present study, but allows us to directly interrogate the effects of choroidal volume change.

^{24,51}This indicates that the mechanical properties of the Bruch's membrane may be important, and a better understanding of how the Bruch's membrane and choroid interact biomechanically will improve future FE models of the eye. In addition, Bruch's membrane is a complex tissue with five sublayers, which may impact its biomechanical behavior; thus, our assumption of a linear-elastic and homogenous material is likely an oversimplification.

**A.J. Feola**, None;

**E.S. Nelson**, None;

**J. Myers**, None;

**C.R. Ethier**, None;

**B.C. Samuels**, None

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