Optimization design method and system of tri-leaflet artificial biological aortic valve leaflets

Abstract

The application discloses an optimization design method and system of a tri-leaflet artificial biological aortic valve leaflet, and the optimization design method comprises the following steps: constructing a tri-leaflet artificial biological aortic valve leaflet geometric model, wherein the tri-leaflet artificial biological aortic valve leaflet geometric model takes a leaflet free edge fillet radius and a leaflet thickness as variable geometric parameters; on the basis of the tri-leaflet artificial biological aortic valve leaflet geometric model, a plurality of different leaflet free edge fillet radii are given, and a plurality of first candidate leaflet models are generated; multi-index simulation evaluation is carried out; according to the multi-index simulation evaluation results of the plurality of first candidate leaflet models, a target leaflet free edge fillet radius is determined; on the basis of the tri-leaflet artificial biological aortic valve leaflet geometric model with the target leaflet free edge fillet radius, a plurality of different leaflet thicknesses are given, and a plurality of second candidate leaflet models are generated; multi-index simulation evaluation is carried out; and according to the multi-index simulation evaluation results, a target leaflet thickness is determined.

Description

Technical Field

[0001] This application relates to the field of artificial biological heart valve design technology, and in particular to an optimized design method, system, computer device, computer-readable storage medium, and computer program product for a tricuspid artificial biological aortic valve leaflet. Background Technology

[0002] Artificial aortic valves typically achieve unidirectional blood flow through three flexible leaflets that open and close during the cardiac cycle. The geometry of the leaflets directly affects the valve opening area, the stress state of the leaflets, and the flow field distribution in the aorta behind the valve. Specifically, the free edge region and the ventral region of the leaflets are prone to significant deformation and localized stress concentration during opening and closing; the leaflet thickness affects the overall flexibility, opening capacity, and structural load-bearing capacity of the leaflets. Therefore, the radius of the free edge radius and the leaflet thickness are crucial geometric parameters in the structural design of artificial aortic valves.

[0003] Current methods for designing artificial heart valves typically involve constructing leaflet models based on empirical geometric models or a few pre-set parameters, followed by evaluation of valve performance using finite element analysis, fluid dynamics analysis, or fluid-structure interaction analysis. While these methods can reflect the valve's structural mechanical response or hemodynamic performance to some extent, in actual design, they often focus more on a single evaluation metric, such as maximum stress, opening area, or flow field distribution, lacking a phased, attributable optimization process for leaflet geometric parameters. Furthermore, when multiple geometric parameters change simultaneously, it is difficult to determine the specific impact of a particular parameter on valve opening performance, leaflet stress-strain levels, and aortic flow field parameters.

[0004] Especially for tricuspid aortic valves, the effects of the free edge radius and leaflet thickness on valve performance are not simply linear. An excessively small free edge radius can lead to stress concentration near the free edge, while variations in the radius can affect the valve's opening morphology and postvalvular flow field. While excessively thin leaflets are beneficial for opening, they may result in greater deformation and stress-strain levels, while excessively thick leaflets may increase structural stiffness but restrict valve opening. Selecting leaflet geometry parameters based on a single metric may make it difficult to achieve a reasonable balance between opening performance, structural stability, and hemodynamic performance.

[0005] Therefore, existing technologies still require a geometric parameter optimization design method for the leaflets of tricuspid artificial aortic valves. This method should use the leaflet free edge radius and leaflet thickness as variable geometric parameters, and determine the target leaflet free edge radius and target leaflet thickness in stages through simulation evaluation of multiple candidate models. This would provide a more explicit, repeatable parameter determination scheme for the design of tricuspid artificial aortic valve leaflets that takes into account opening performance, structural mechanical performance, and hemodynamic performance. Summary of the Invention

[0006] In view of the shortcomings of the above-mentioned related technologies, the purpose of this application is to provide an optimized design method, system, computer device, computer-readable storage medium, and computer program product for a trilobal artificial biological aortic valve leaflet.

[0007] To achieve the above and other related objectives, the first aspect of this application provides an optimization design method for the leaflets of a tricuspid artificial aortic valve. The optimization design method includes the following steps: constructing a geometric model of the leaflets of the tricuspid artificial aortic valve, wherein the leaflet geometric model uses the leaflet free edge radius and leaflet thickness as variable geometric parameters; based on the leaflet geometric model, assigning multiple different leaflet free edge radii to generate multiple first candidate leaflet models; and performing multi-index simulation evaluation on the multiple first candidate leaflet models, wherein the multi-index simulation evaluation includes an opening performance evaluation and a structural mechanical evaluation based on the dynamic opening and closing process of the leaflets under transvalvular pressure differential. The evaluation includes at least two types of evaluations, such as price and hemodynamic evaluation based on the aortic flow field model; determining the target leaflet free edge radius based on the multi-index simulation evaluation results of multiple first candidate leaflet models; generating multiple second candidate leaflet models by assigning multiple different leaflet thicknesses based on the leaflet geometric model with the target leaflet free edge radius; performing the multi-index simulation evaluation on the multiple second candidate leaflet models; and determining the target leaflet thickness based on the multi-index simulation evaluation results of the multiple second candidate leaflet models to obtain the target leaflet free edge radius and target leaflet thickness for the design of the tricuspid artificial biological aortic valve leaflet.

[0008] A second aspect of this application provides an optimization design system for the leaflets of a tricuspid artificial aortic valve, comprising: a leaflet geometry model construction module for constructing a geometric model of the leaflets of the tricuspid artificial aortic valve, wherein the leaflet geometry model uses the leaflet free edge radius and leaflet thickness as variable geometric parameters; a first candidate model generation module for generating multiple first candidate leaflet models by assigning multiple different leaflet free edge radii to the leaflet geometry model; and a first simulation evaluation module for performing multi-index simulation evaluation on the multiple first candidate leaflet models, wherein the multi-index simulation evaluation includes ostiole performance evaluation, structural mechanics evaluation based on the dynamic opening and closing process of the leaflets under transvalvular pressure differential, and hemodynamics based on an aortic flow field model. The evaluation includes at least two types of assessments: a target fillet radius determination module, used to determine the target leaflet free edge fillet radius based on the multi-index simulation evaluation results of multiple first candidate leaflet models; a second candidate model generation module, used to generate multiple second candidate leaflet models by assigning multiple different leaflet thicknesses to leaflet geometric models with the target leaflet free edge fillet radius; a second simulation evaluation module, used to perform the multi-index simulation evaluation on the multiple second candidate leaflet models; and a target thickness determination module, used to determine the target leaflet thickness based on the multi-index simulation evaluation results of the multiple second candidate leaflet models, so as to obtain the target leaflet free edge fillet radius and target leaflet thickness for the design of the tricuspid artificial biological aortic valve leaflet.

[0009] A third aspect of this application provides a computer device, comprising: a storage device for storing at least one program; and a processing device connected to the storage device, the processing device being configured to call and execute the at least one program from the storage device to implement the optimized design method for the leaflets of a trilobed artificial biological aortic valve as described in the first aspect of this application.

[0010] A fourth aspect of this application provides a computer-readable storage medium storing at least one program that, when executed by a processor, implements the optimized design method for the leaflets of a trilobed artificial biological aortic valve as described in the first aspect of this application.

[0011] The fifth aspect of this application provides a computer program product, which includes a computer program or instructions that, when executed by a processor, implement the optimized design method for the leaflets of the trilobed artificial biological aortic valve as described in the first aspect of this application.

[0012] In summary, the optimization design method, system, computer equipment, computer-readable storage medium, and computer program product for the tricuspid artificial aortic valve leaflets provided in this application use the leaflet free edge radius and leaflet thickness as two independent geometric design variables. First, while keeping the leaflet thickness constant, multiple candidate leaflet models corresponding to different leaflet free edge radii are compared to determine the target leaflet free edge radius. Then, while keeping the target leaflet free edge radius unchanged, multiple candidate leaflet models corresponding to different leaflet thicknesses are compared to determine the target leaflet thickness. Compared to optimizing multiple geometric parameters simultaneously, this application employs a staged, single-variable optimization strategy, which effectively reduces the coupling effect between different design variables. This allows performance differences between different candidate models to be primarily attributed to changes in the corresponding design variables, thereby improving the reliability and interpretability of the parameter optimization results.

[0013] Furthermore, this application not only evaluates the geometric opening area of ​​the valve in the open state, but also combines the stress, strain, blood flow velocity, wall pressure, and wall shear stress in the aortic flow field during the dynamic opening and closing process of the valve leaflet under the action of transvalvular pressure difference to comprehensively evaluate different candidate models. This avoids selecting design parameters based solely on a single evaluation index, thereby taking into account valve opening performance, structural stability, and hemodynamic performance, and improving the rationality of the target leaflet free edge radius and target leaflet thickness.

[0014] This application obtains material parameters based on uniaxial tensile experiments of bovine pericardium and uses a hyperelastic constitutive model to characterize the valve leaflet material, enabling the valve leaflet material to more accurately reflect the large deformation nonlinear mechanical response of the artificial biological aortic valve during the cardiac cycle. Compared with using empirical material parameters or linear elastic material models, this application can improve the consistency between structural mechanics simulation analysis and actual working conditions, thereby improving the credibility of parameter optimization results.

[0015] This application, based on the comprehensive evaluation results of multiple candidate leaflet models, further screens candidate models that meet the preset opening performance conditions by combining structural mechanics evaluation results and hemodynamic evaluation results to determine the target leaflet free edge radius and target leaflet thickness. By establishing a unified evaluation process and parameter screening rules, the determination process of leaflet geometric parameters has a clear design basis, reducing the uncertainty brought about by empirical design and improving the standardization, repeatability, and engineering application value of the artificial biological aortic valve leaflet optimization design process. Attached Figure Description

[0016] The specific features involved in this application are shown in the appended claims. The features and advantages of the invention can be better understood by referring to the exemplary embodiments and accompanying drawings described in detail below. A brief description of the drawings is as follows:

[0017] Figure 1 The flowchart shown is a method for optimizing the design of the leaflets of a trilobed artificial biological aortic valve in one embodiment of this application.

[0018] Figure 2 The diagram shown is a schematic representation of the region division of a basic leaflet unit in one embodiment of this application.

[0019] Figure 3 The diagram shown is a geometric model of a basic leaflet unit in one embodiment of this application.

[0020] Figure 4 The diagram shown is a structural schematic of a trilobate artificial biological aortic valve in one embodiment of this application.

[0021] Figure 5 The diagram shown is a schematic representation of the uniaxial tensile stress-strain curve of a valve in one embodiment of this application.

[0022] Figure 6 The diagram shown is a schematic representation of the inlet and outlet pressures of a model in one embodiment of this application.

[0023] Figure 7 The image shown is a top view of six sets of trilobed artificial biological aortic valves with different leaflet free edge radii in one embodiment of this application.

[0024] Figure 8 The diagram shows the stress distribution cloud map of the valve in the fully open and fully closed states under different grid densities in one embodiment of this application.

[0025] Figure 9 This is a schematic diagram illustrating the dynamic change of valve displacement during a cardiac cycle in one embodiment of this application.

[0026] Figure 10 The diagram shown illustrates the calculation of the valve geometric opening area in one embodiment of this application.

[0027] Figure 11a The graph shown is a curve illustrating the change of the maximum principal stress on the valve over time in one embodiment of this application.

[0028] Figure 11b The graph shown is a curve illustrating the change of the average maximum principal stress on the valve over time in one embodiment of this application.

[0029] Figure 12The diagram shows the maximum principal stress cloud map of the first candidate leaflet model at different times in one embodiment of this application.

[0030] Figure 13a The graph shown is a curve of the maximum principal strain on the valve versus time in one embodiment of this application.

[0031] Figure 13b The graph shown is a curve of the average maximum principal strain versus time in one embodiment of this application.

[0032] Figure 14 The image shown is a contour plot of the maximum principal strain on the valve at different times for a first candidate leaflet model in one embodiment of this application.

[0033] Figure 15a The graph shown is a curve of the maximum equivalent stress on the valve over time in one embodiment of this application.

[0034] Figure 15b The graph shown is a curve illustrating the change of average maximum equivalent stress over time in one embodiment of this application.

[0035] Figure 16 The diagram shows the equivalent stress cloud at different times for the first candidate leaflet model in one embodiment of this application.

[0036] Figure 17a The graph shown is a curve illustrating the change of maximum shear stress on the valve over time in one embodiment of this application.

[0037] Figure 17b The graph shown is a curve illustrating the change of the average maximum shear stress on the valve over time in one embodiment of this application.

[0038] Figure 18 The diagram shows the maximum shear stress cloud map of the first candidate leaflet model at different times in one embodiment of this application.

[0039] Figure 19 The image shown is a leaflet geometry model at peak systolic aortic valve position in one embodiment of this application.

[0040] Figure 20 This is a schematic diagram showing the typical planar position of the aorta in one embodiment of this application.

[0041] Figure 21 The image shown is a cloud map of blood flow patterns of valves with different fillet radii during the peak systolic phase in the same aortic flow field, as presented in one embodiment of this application.

[0042] Figure 22 This diagram illustrates the maximum blood flow velocity of valves with different fillet radii during the peak systolic phase in the same aortic flow field, as shown in one embodiment of this application.

[0043] Figure 23a The diagram shows the maximum blood flow velocity in a segmented aortic region in one embodiment of this application.

[0044] Figure 23b The diagram shows the average blood flow velocity in the aortic region in one embodiment of this application.

[0045] Figure 24 The image shown is a cross-sectional cloud diagram of blood flow velocity in the same aorta at different corner radii during the peak systolic period in one embodiment of this application.

[0046] Figure 25 The image shown is a cloud map of vascular pressure distribution at peak systolic time for valves with different corner radii in one embodiment of this application.

[0047] Figure 26 This diagram illustrates the maximum wall pressure at the peak systolic time for valves with different corner radii in one embodiment of this application.

[0048] Figure 27 The diagram shows the average wall pressure of each model in the aortic segment region in one embodiment of this application.

[0049] Figure 28 The image shown is a cloud map illustrating the wall shear stress distribution of heart valves with different fillet radii during peak systole within the same aorta in one embodiment of this application.

[0050] Figure 29 The diagram shows the maximum wall shear stress at the peak systolic time for valves with different corner radii in one embodiment of this application.

[0051] Figure 30 The diagram shows the average wall shear stress of each model in the aortic segment region in one embodiment of this application.

[0052] Figure 31 The image shown is a top view of five sets of trilobed artificial biological aortic valves with different leaflet thicknesses in one embodiment of this application.

[0053] Figure 32 The diagram shows the dynamic changes of the second candidate leaflet model on the valve at different times in one embodiment of this application.

[0054] Figure 33a The graph shown is a curve illustrating the change of maximum principal stress on the valve over time in one embodiment of this application.

[0055] Figure 33b The graph shown is a curve illustrating the change of the average maximum principal stress on the valve over time in one embodiment of this application.

[0056] Figure 34 The diagram shows the maximum principal stress cloud map on the valve at different times for five groups of second candidate leaflet models in one embodiment of this application.

[0057] Figure 35a The graph shown is a curve of the maximum principal strain on the valve versus time in one embodiment of this application.

[0058] Figure 35b The graph shown is a curve of the average maximum principal strain versus time in one embodiment of this application.

[0059] Figure 36 The image shown is a contour plot of the maximum principal strain on the valve at different times for five groups of second candidate leaflet models in one embodiment of this application.

[0060] Figure 37a The graph shown is a curve illustrating the change of maximum equivalent stress on the valve over time in one embodiment of this application.

[0061] Figure 37b The graph shown is a curve illustrating the change of average maximum equivalent stress over time in one embodiment of this application.

[0062] Figure 38 The diagram shows the equivalent stress cloud map of the valve at different times for five groups of second candidate leaflet models in one embodiment of this application.

[0063] Figure 39a The graph shown is a curve illustrating the change of maximum shear stress on the valve over time in one embodiment of this application.

[0064] Figure 39b The graph shown is a curve illustrating the change of the average maximum shear stress on the valve over time in one embodiment of this application.

[0065] Figure 40 The diagram shows the maximum shear stress cloud map of the five groups of second candidate leaflet models at different times in one embodiment of this application.

[0066] Figure 41 The image shown is a cloud map of the blood velocity streamline distribution of each thickness model during the peak contraction period in one embodiment of this application.

[0067] Figure 42 This is a schematic diagram showing the maximum blood velocity of each thickness model during the peak contraction period in one embodiment of this application.

[0068] Figure 43a The diagram shows the maximum blood flow velocity in a segmented aortic region in one embodiment of this application.

[0069] Figure 43b The diagram shows the average blood flow velocity in the aortic region in one embodiment of this application.

[0070] Figure 44 The image shown is a cross-sectional cloud diagram of the intraluminal velocity of the aorta during peak systole, as described in one embodiment of this application.

[0071] Figure 45a The image shown is a cloud map of the pressure distribution on the aortic wall under the same conditions for models of different thicknesses in one embodiment of this application.

[0072] Figure 45b The figure shows the maximum wall pressure under the same conditions for different thickness models in one embodiment of this application.

[0073] Figure 46 The diagram shown is a schematic representation of the average wall pressure in a segmented aortic region according to one embodiment of this application.

[0074] Figure 47a The image shown is a cloud map of the shear stress distribution on the aortic wall under the same conditions for models of different thicknesses in one embodiment of this application.

[0075] Figure 47b The figure shows the maximum wall shear stress under the same conditions for different thickness models in one embodiment of this application.

[0076] Figure 48 The diagram shown is a schematic representation of the average wall shear stress in a segmented region of the aorta in one embodiment of this application.

[0077] Figure 49 The diagram shown is a block diagram of an optimized design system for a trilobed artificial biological aortic valve leaflet in one embodiment of this application.

[0078] Figure 50 The diagram shown is a structural schematic of a computer device according to one embodiment of this application. Detailed Implementation

[0079] The following specific embodiments illustrate the implementation of this application. Those skilled in the art can easily understand the advantages and technical effects of this application from the content disclosed in this specification. In the following description, some embodiments may refer to the accompanying drawings. It should be understood that other embodiments not shown in the drawings may also be used, and specific steps, modules or units, electrical and operational changes may be made without departing from the spirit and scope of this application. The detailed description below should not be considered limiting, and the scope of the embodiments of this application is limited only by the claims published in this application. The terminology used herein is for describing particular embodiments only and is not intended to limit this application.

[0080] As used herein, the terms "computer," "personal computer," and "computing device" refer to any programmable computer system known or to be developed in the future. In a preferred embodiment, the computer will be coupled to a network such as that described herein. The computer system may be configured with processor-executable software instructions to perform the processes described herein.

[0081] As used herein, the terms “module” and “system” are intended to refer to a computer-related entity, whether that entity is hardware, a combination of hardware and software, software, or software in execution. For example, a component can be, but is not limited to, a process running on a processor, a processor, an object, an executable, a thread in execution, a program, and / or a computer. In the example, both an application running on a server and the server itself can be components. One or more components may reside within a process and / or a thread in execution, and components may be localized on a single computer and / or distributed across two or more computers.

[0082] In this application, unless otherwise stated, the terms "first," "second," etc., are used only to distinguish different stages, different objects, or different parameters, and do not indicate any limitation on quantity, order, or importance. For example, the first candidate leaflet model is used to evaluate the effect of changes in the radius of the leaflet's free edge fillet, and the second candidate leaflet model is used to evaluate the effect of changes in the leaflet's thickness.

[0083] As described in the background section, existing technologies have proposed methods for parametric modeling of artificial heart valves using variable parametric geometric curves, combined with finite element simulation, multi-objective optimization algorithms, and parameter combination screening methods to optimize the valve geometry. While these methods can improve the automation and design freedom of valve geometry design, they typically focus on optimizing the overall geometry of the artificial heart valve or high-degree-of-freedom parameters controlled by multiple curves, easily leading to a high-dimensional parameter search space. Within this optimization framework, the effects of different geometric parameters on the valve opening area, leaflet stress-strain distribution, and the aortic flow field behind the valve are often coupled, making it difficult to clearly determine the specific contribution of a key local geometric parameter to valve performance changes. In particular, these methods do not establish a phased optimization path for the two key parameters of the tricuspid aortic valve leaflet free edge radius and leaflet thickness—first determining the free edge radius and then determining the leaflet thickness based on the target free edge radius—thus failing to provide a clear attribution basis for parameter determination in the local geometric design of the artificial aortic valve leaflets.

[0084] Before describing the specific implementation methods of this application, some basic terms involved in this application will be explained. It should be noted that the following terminology descriptions are used to help understand the technical solutions of this application and should not be construed as additional limitations on the scope of protection of this application; where there is no contradiction, the following terminology descriptions can be used in conjunction with the specific technical features of each embodiment.

[0085] In this application, a tricuspid artificial aortic valve refers to an artificial biological valve structure used to simulate, design, or replace the function of a natural aortic valve. It comprises three flexible leaflets capable of opening and closing during the cardiac cycle. The three leaflets are arranged circumferentially around the valve's central axis, forming an orifice for blood flow when open, and cooperating to prevent backflow when closed. The leaflets of the tricuspid artificial aortic valve can be made of bovine pericardium, porcine pericardium, porcine aortic valve tissue, or other biomaterials suitable for artificial biological valves. In some specific embodiments, the leaflets are made of bovine pericardium.

[0086] In this application, the leaflet refers to a flexible, sheet-like structure in a tricuspid aortic valve that can deform, open, and close under transvalvular pressure differential. The leaflet may include a fixed edge for connection to the valve annulus, stent, or other fixation structures, a ventral curved surface located in the main body region of the leaflet, and a free edge for mating or apposition with adjacent leaflets when the valve closes. This application primarily focuses on optimizing the local geometric and thickness parameters of the leaflet, without requiring limitations on the stent structure, suture structure, assembly structure, or implantation structure of the aortic valve.

[0087] In this application, the leaflet geometric model refers to a computational model used to characterize the spatial morphology of the leaflets of a tricuspid artificial aortic valve. This leaflet geometric model can be a surface model, a solid model, a CAD model, a finite element preprocessing model, or other geometric models that can be recognized by simulation software and used for structural mechanics analysis and / or hemodynamic analysis. The leaflet geometric model uses the leaflet free edge fillet radius and leaflet thickness as variable geometric parameters, and different candidate leaflet models can be generated by changing the leaflet free edge fillet radius or the leaflet thickness.

[0088] In this application, the leaflet free edge fillet radius refers to a geometric parameter used to characterize the degree of local arc transition at the leaflet free edge. In some embodiments, the leaflet free edge fillet radius is the radius of the arc used to form the leaflet free edge. A larger leaflet free edge fillet radius indicates a smoother arc transition at the leaflet free edge; a smaller leaflet free edge fillet radius indicates a steeper arc transition at the leaflet free edge. By changing the leaflet free edge fillet radius, multiple first candidate leaflet models with different local transition morphologies can be generated to evaluate the impact of changes in the free edge fillet radius on leaflet opening performance, structural mechanical properties, and / or hemodynamic performance.

[0089] In this application, leaflet thickness refers to the thickness parameter of the leaflet in a direction perpendicular to or approximately perpendicular to the mid-surface of the leaflet. In some embodiments, the leaflet thickness can be a uniform thickness within the leaflet region; in other embodiments, the leaflet thickness can also be an equivalent thickness or average thickness used to characterize the overall thickness level of the leaflet. By changing the leaflet thickness, multiple second candidate leaflet models with different thicknesses can be generated to evaluate the effects of leaflet thickness variations on leaflet opening performance, structural mechanical properties, and / or hemodynamic performance.

[0090] In this application, a candidate leaflet model refers to a leaflet model generated by assigning different geometric parameters to a leaflet geometric model, and used for simulation evaluation and parameter selection. The candidate leaflet models include a first candidate leaflet model and a second candidate leaflet model. The first candidate leaflet model is a candidate model generated by assigning different leaflet free edge fillet radii; multiple first candidate leaflet models are used to determine the target leaflet free edge fillet radius. The second candidate leaflet model is a candidate model generated by assigning different leaflet thicknesses based on a determined target leaflet free edge fillet radius; multiple second candidate leaflet models are used to determine the target leaflet thickness. Furthermore, the same candidate leaflet model can have different data representations at different processing stages. For example, in the geometric modeling stage, it may be represented as a candidate leaflet geometric model; in the structural mechanics evaluation stage, it may be represented as a corresponding candidate leaflet finite element model, where the finite element model is a structural mechanics calculation model formed after finite element preprocessing based on the geometric model; in the hemodynamics evaluation stage, it may be represented as an aortic flow field model based on its maximum opening configuration. During application, the appropriate data representation format can be adopted according to data requirements.

[0091] In this application, the target leaflet free edge radius refers to the free edge radius determined from multiple different leaflet free edge radii based on the multi-index simulation evaluation results of multiple first candidate leaflet models, and is used for the design of the tricuspid artificial biological aortic valve leaflet. The target leaflet thickness refers to the thickness parameter determined from multiple different leaflet thicknesses based on the multi-index simulation evaluation results of multiple second candidate leaflet models, and is used for the design of the tricuspid artificial biological aortic valve leaflet, given that the target leaflet free edge radius has already been determined.

[0092] In this application, multi-index simulation evaluation refers to the process of performing simulation analysis on candidate leaflet models and comparing and selecting candidate leaflet models based on at least two types of evaluation results. The at least two types of evaluation are selected from opening performance evaluation, structural mechanics evaluation, and hemodynamic evaluation. In some embodiments, multi-index simulation evaluation simultaneously includes opening performance evaluation, structural mechanics evaluation, and hemodynamic evaluation. Specifically, opening performance evaluation characterizes the ability of the candidate leaflet model to form a blood flow channel in the open state; structural mechanics evaluation characterizes the stress, strain, or other structural responses of the candidate leaflet model during dynamic opening and closing under transvalvular pressure differential; and hemodynamic evaluation characterizes the impact of the candidate leaflet model on flow field indicators such as blood flow velocity, wall pressure, and / or wall shear stress after being implanted into the aortic flow field model.

[0093] Please see Figure 1 The figure shows a flowchart of an optimization design method for a trilobed artificial biological aortic valve leaflet in one embodiment of this application. As shown in the figure, the optimization design method includes steps S100 to S160.

[0094] In step S100, a leaflet geometric model of a trilobate artificial aortic valve is constructed, wherein the leaflet geometric model uses the leaflet free edge radius and leaflet thickness as variable geometric parameters.

[0095] In one embodiment, the leaflet geometry model includes a basic leaflet unit. The basic leaflet unit refers to the fundamental geometric unit used to form a single leaflet in a tricuspid aortic valve. The three leaflets are formed by circumferentially arranging the basic leaflet unit around the central axis of the tricuspid aortic valve. That is, a basic leaflet unit can be circumferentially arranged around the central axis of the tricuspid aortic valve at a predetermined angle to form three leaflets, thereby obtaining a complete tricuspid leaflet geometry model.

[0096] Please see Figure 2The figure shows a schematic diagram of the region division of a basic leaflet unit in one embodiment of this application. As shown, the basic leaflet unit includes a fixed edge 11, a ventral surface 12, and a free edge 13. The fixed edge 11 refers to the boundary region in the basic leaflet unit used to represent the connection between the leaflet and the valve annulus, stent, suture edge, or other fixed structures. In finite element dynamic opening and closing analysis, the fixed edge 11 can serve as a region for applying fixed constraints, displacement constraints, or rotational constraints to simulate the fixed connection state of the leaflet in an artificial biological aortic valve. The fixed edge 11 differs from the free edge 13, and it is generally not used as a free deformation boundary during the dynamic opening and closing of the leaflet. The ventral surface 12 refers to the main curved surface region of the leaflet located between the fixed edge 11 and the free edge 13. The ventral surface 12 is the main region where the leaflet undergoes bending, stretching, shearing, and opening and closing deformation under the action of transvalvular pressure differential. In some embodiments, the ventral surface 12 can be generated by rotating the arc forming the free edge 13 about a predetermined axis, thereby forming a composite surface with radial spherical curvature and circumferential arc curvature. The abdominal curved surface 12 can also be understood as the curved surface area used to connect the fixed edge 11 and the free edge 13, and to determine the three-dimensional shape of the leaflet body.

[0097] In one embodiment, constructing the geometric model of the basic leaflet unit includes: using two opposing boundary lines as the boundaries of the basic leaflet unit, and forming the free edge by connecting the two boundary lines with a circular arc, the radius of which is the fillet radius of the leaflet free edge.

[0098] The boundary line refers to the geometric line segment or curve used to define the lateral boundary of a single leaflet when constructing a basic leaflet unit. Two opposing boundary lines, together with an arc connecting the two boundary lines, define the free edge of the basic leaflet unit and its adjacent region. In some specific embodiments, the two boundary lines may correspond to two tangents in the leaflet geometry model, and the arc may connect the two tangents to form the leaflet free edge.

[0099] The arc is preferably a continuous, smooth curve, with a smooth transition between the arc and the two boundary lines to avoid geometrical abrupt changes at the connection between the free edge and the side boundary. The radius of the arc is defined as the fillet radius of the leaflet free edge, which serves as one of the parametric design variables of the basic leaflet unit and characterizes the degree of curvature of the free edge. When the fillet radius increases, the curvature of the free edge decreases, and the free edge profile tends to be flatter; when the fillet radius decreases, the curvature of the free edge increases, and the free edge profile becomes more curved.

[0100] Furthermore, during parametric modeling, only the radius of the arc needs to be adjusted to generate basic leaflet units with different free edge morphologies, while the positional relationship of the two boundary lines, the position of the fixed edge, and the geometric topology of the basic leaflet units remain unchanged. In other words, different candidate models maintain the same geometric topology, with only the free edge curvature changing, thus ensuring a consistent geometric basis and comparability among different candidate models. This allows the evaluation results obtained from subsequent structural mechanics and hemodynamics analyses to accurately reflect the impact of changes in the free edge radius of the leaflet.

[0101] Preferably, the arcs are tangentially connected to the two boundary lines respectively, so that a continuous first derivative curve is formed between the free edge and the side boundary, thereby improving the continuity and smoothness of the leaflet geometric model, reducing the possibility of local geometric abrupt changes at the free edge connection position, and reducing local stress concentration caused by discontinuity of the model boundary.

[0102] After the free edge is constructed, it serves as the basic contour for generating the ventral surface of the leaflet, and is further combined with the fixed edge and two boundary lines to form a complete basic leaflet unit. Subsequently, a complete leaflet geometric model of the tricuspid artificial aortic valve can be established based on the basic leaflet unit, and the fillet radius of the leaflet free edge is used as the first geometric parameter in the subsequent optimization design stage.

[0103] By employing the aforementioned modeling method, the basic outline of the leaflet is defined by two boundary lines, and free edges are formed by connecting them with circular arcs. On the one hand, this approach can accurately describe the basic geometric characteristics of the leaflet of the tricuspid artificial aortic valve with fewer geometric parameters, making the free edge radius an independent design variable with clear physical meaning. On the other hand, while maintaining the overall geometric topology of the leaflet, multiple candidate leaflet models can be quickly generated simply by changing the radius of the circular arc. This facilitates subsequent evaluations of opening performance, structural mechanics, and hemodynamics under different free edge radius conditions, and improves the repeatability and attributability of simulation results among different candidate models.

[0104] In one embodiment, constructing the geometric model of the basic leaflet unit further includes rotating the arc forming the free edge (also called the free side) about an axis to generate the ventral surface of the basic leaflet unit; wherein the axis is a line connecting the ends of the two boundary lines that are away from the arc.

[0105] In this application, the axis refers to the geometric reference line for rotating the arc when constructing the ventral surface of the basic leaflet unit. In some embodiments, the axis is a line connecting the ends of two boundary lines that are away from the arc. The ventral surface of the basic leaflet unit can be generated by rotating the arc forming the free edge about the axis by a predetermined angle.

[0106] Specifically, a line is formed by connecting the ends of the two boundary lines furthest from the free edge, and this line is defined as the axis of rotation. Subsequently, the arc forming the free edge is rotated about the axis of rotation to generate the ventral surface of the basic leaflet unit. During rotation, each point on the arc moves in a circle about the axis of rotation, and their trajectories collectively form a continuous spatial surface, thus constituting the ventral surface of the basic leaflet unit.

[0107] In this embodiment, the rotation axis is located on the side of the basic leaflet unit near the fixed end of the leaflet and passes through the end of the two boundary lines away from the free edge. Since the free edge maintains a preset distance from the rotation axis, after the free edge rotates around the rotation axis, it can form a continuous curved surface that gradually transitions from the fixed edge to the free edge, so that the ventral curved surface has a continuously changing curvature in both the radial and circumferential directions.

[0108] Furthermore, during rotational modeling, the arc forming the free edge maintains its own geometric shape and only rotates spatially around the rotation axis. The resulting ventral surface inherits the geometric features of the free edge curve. When the radius of the free edge fillet is adjusted, the curvature of the free edge arc changes, and the generated ventral surface is automatically updated accordingly. This establishes a parametric relationship between the free edge fillet radius and the overall surface morphology of the leaflet. In other words, the free edge fillet radius not only determines the free edge contour but also synchronously influences the spatial geometry of the ventral surface of the leaflet through rotational modeling, ensuring consistent modeling methods and geometric topology across different candidate leaflet models.

[0109] Preferably, the abdominal curved surface is generated by rotating the free edge arc continuously in one step, so that the entire leaflet surface remains continuous and smooth, and there are no broken line transition areas or curvature abrupt change areas formed by splicing multiple curved surfaces, thereby improving the continuity and smoothness of the leaflet geometric model.

[0110] Furthermore, after generating the abdominal surface, a closed leaflet geometric contour can be formed by combining the fixed edge, two boundary lines, and free edge, and a preset leaflet thickness can be assigned to establish a complete basic leaflet unit. Subsequently, the basic leaflet unit is circumferentially arrayed around the central axis of the tricuspid artificial aortic valve to form a complete tricuspid artificial aortic valve leaflet geometric model.

[0111] By using the above modeling method, the ventral surface of the leaflet is generated by rotating the free edge arc around a preset rotation axis. This not only forms a continuous surface that conforms to the spatial morphology of the trilobed artificial aortic valve, but also establishes a parametric relationship between the free edge fillet radius and the overall geometric shape of the leaflet. This allows different candidate leaflet models to generate the corresponding leaflet surface simply by adjusting the free edge fillet radius, while maintaining the same geometric topology and modeling rules. This provides a unified and repeatable parametric geometric model for subsequent optimization of the free edge fillet radius and leaflet thickness.

[0112] In one embodiment, the arc is rotated about the axis by an angle of 120°, so that the basic leaflet unit forms a composite surface with radial spherical curvature and circumferential arc curvature.

[0113] After modeling a single basic leaflet unit, the basic leaflet unit is then used as an array unit and circumferentially replicated around the central axis of the tricuspid artificial aortic valve at a preset angle to form three leaflets evenly distributed circumferentially, thus establishing a complete leaflet geometric model of the tricuspid artificial aortic valve. The leaflet array replication step is independent of the step of generating the abdominal surface from the free-edge arc around the rotation axis; the former is used to construct the spatial surface of a single leaflet unit, while the latter is used to construct the complete leaflet geometric model of the tricuspid artificial aortic valve.

[0114] Please see Figure 3 The figure shows a schematic diagram of the geometric model of the basic leaflet unit in one embodiment of this application. As shown in the figure, the construction process of the geometric model of the basic leaflet unit includes: taking two tangents SM and TN as boundaries, and connecting them in the middle through an arc MN to form the free edge of the basic leaflet unit, with the center of the arc being Or and the radius being r; rotating the arc around the ST vector by 120° to generate the surface of the basic leaflet unit, forming a composite geometric shape with radial spherical curvature and circumferential arc curvature.

[0115] exist Figure 3 Based on the geometric model of the basic leaflet unit shown, three symmetrical basic leaflet units are generated by circular array with the y-axis as the rotation axis, thereby obtaining the complete leaflet geometric model of the tricuspid artificial aortic valve.

[0116] In one specific embodiment, the leaflet geometry model can be constructed using the 3D modeling software SolidWorks. See also... Figure 4 The figure shows a schematic diagram of the structure of a trilobate artificial biological aortic valve in one embodiment of this application. The figure shows the top view, front view and left view of the trilobate artificial biological aortic valve.

[0117] Based on the aforementioned leaflet geometry model, finite element preprocessing is performed on the leaflet geometry model to construct a finite element model of the tricuspid artificial aortic valve for dynamic opening and closing analysis. The finite element preprocessing includes assigning material properties to the leaflets, generating a finite element mesh, setting the contact relationship between the three leaflets, setting the analysis step, applying transvalvular pressure differential load, and setting boundary constraints.

[0118] In one embodiment, since uniaxial tensile fracture testing is a key method for evaluating the mechanical properties of valve materials, it can quantify the material's ultimate strength and elongation at break, reflecting its fracture resistance and deformation characteristics, and providing preliminary evidence for the material's durability during the heart's periodic motion. Furthermore, the stress-strain curves plotted based on experimental data can lay the foundation for subsequent establishment of a valve constitutive model. Therefore, in one embodiment, the material parameters of the valve leaflet geometry model are obtained based on uniaxial tensile tests of bovine pericardium material. Specifically, bovine pericardium material is used as the valve leaflet material, and the constitutive model parameters of the valve leaflet material are obtained through uniaxial tensile testing.

[0119] In one example, a pericarp sample was cut into 2×35 mm uniaxial tensile standard pieces using a dumbbell-shaped die. The load on the unloaded tensile testing machine was zeroed, and the sample was horizontally clamped at both ends of the tensile testing machine to prevent slippage. The clamps at both ends were operated to slowly stretch the sample until the sensor received a signal of tensile force of 0.1 N, which served as a sign that the sample had been flattened and the tensile test had begun. A constant tensile rate of 0.01 mm / s was set, and a quasi-static uniaxial tensile test was performed, recording the load and displacement in real time until the material broke. After the test, the tensile force F in the tensile axis direction and the elongation L relative to the initial state were obtained. The nominal stress σ and nominal strain ε of the sample were calculated according to formulas (1) and (2), respectively:

[0120] Formula (1)

[0121] Formula (2)

[0122] Where A is the initial cross-sectional area of ​​the specimen, which can be determined based on the initial width and initial thickness of the specimen; L is the elongation of the specimen in the tensile direction, and L0 is the initial gauge length of the specimen. The initial thickness of the specimen is measured before the tensile test, and the initial cross-sectional area A is determined based on the product of the initial width and initial thickness of the specimen.

[0123] Please see Figure 5The image shows a schematic diagram of the uniaxial tensile stress-strain curve of the valve in one embodiment of this application. Based on the stress-strain data of bovine pericardial tissue obtained from the uniaxial tensile experiment, a constitutive model is fitted to obtain the corresponding constitutive model parameters. In one example, the stress-strain data obtained from the experiment is first processed, then the processed data is imported into appropriate software such as Abaqus, and a hyperelastic material model is selected in the corresponding module for parameter fitting to obtain the corresponding constitutive model parameters.

[0124] After obtaining the constitutive model, the method for constructing a finite element model of the tricuspid artificial aortic valve based on the leaflet geometry model includes: 1) Assigning material parameters: importing the obtained geometric model of the basic leaflet unit into finite element software such as Abaqus in solid form and setting the material density to 1100 kg / m³ in the corresponding module. 3 ; Select the hyperelastic constitutive model obtained from the above screening and input the corresponding material parameters. Establish a homogeneous solid section and assign the above material properties to the section. Integrate the three basic leaflet element geometric models that already have material properties into a whole through the "assembly" operation; 2) Mesh generation: Control the size of the mesh elements and generate a mesh for the model. Specifically, use the "Seed" operation in the software to evenly distribute points on the leaflet edges to control the size of the mesh elements. Select the corresponding element type from the solid elements in the element library to generate a mesh for the model; 3) Contact settings: Since the leaflets will come into contact with each other during the entire simulation process, the contact relationship needs to be set. Set "face-to-face contact" between each pair of the three leaflets, and select the penalty function for the contact pair algorithm; 4) Analysis step: Since the cardiac cycle is 0.8 s, the analysis step time is set to 0.8 s. Insert a new analysis step after the initial analysis step and name it "Step-1". Select the analysis step type as dynamic explicit analysis, the analysis step duration as 0.8 s, and set the result output time interval to 0.02 s; 5) Load and Constraints: Use the ventricular-aortic pressure difference curve as the input load for the leaflet. Apply surface loads to the leaflet in the form of pressure in the "Step-1" analysis step, and import them in tabular form under the amplitude module to simulate the load changes during the opening and closing of the leaflet. Apply fixed constraints to the fixed edge of the leaflet to restrict the movement and rotation of the fixed edge.

[0125] Based on this, a new job is created for the model, submitted, and solved. After the solution is completed, the opening and closing configurations, displacements, stresses, strains, and other structural mechanics calculations of the valve leaflets during the cardiac cycle are obtained.

[0126] It should be noted that the “lobes geometric model” in this application is used to characterize the spatial geometric shape of the lobes, the “finite element model” is a structural mechanics calculation model formed after finite element preprocessing based on the lobes geometric model, and the finite element analysis is the process of solving the finite element model and obtaining the structural response results. The three are not to be understood as equivalent.

[0127] Please see Figure 6 The figure shows a schematic diagram of the inlet and outlet pressure curves of the model in one embodiment of this application. As shown in the figure, four key time points within a cardiac cycle were subsequently analyzed: the initial systolic phase (t1=0.1 s), the peak systolic phase (t2=0.2 s), the initial diastolic phase (t3=0.32 s), and the complete closure phase (t4=0.42 s). This application utilizes the explicit kinetics solver module of Abaqus software to complete the calculations.

[0128] Furthermore, macroscopic observation, histological staining, and scanning electron microscopy analysis of the diseased aortic valve revealed significant thickening, extracellular matrix remodeling, disordered collagen and elastic fiber networks, interlaminar fissures, and lipid and calcification deposits, indicating substantial changes in its material composition and microstructure. This may lead to abnormal local mechanical responses and an increased risk of fatigue damage. For artificial bioprosthetic valves, given the relatively fixed material properties and overall dimensions, local geometry remains a crucial factor influencing valve opening and closing behavior, stress-strain distribution, and structural stability. Existing research indicates that boundary transition forms and local curvature variations significantly affect the formation of high-stress regions. Optimizing the local geometry of the valve is an effective way to improve its mechanical properties and long-term durability. In addition, the mechanical response of artificial bioprosthetic valves is not only affected by boundary transition forms but also closely related to valve thickness. Changes in thickness significantly impact the overall stiffness, deformation capacity, and local stress transmission patterns of the valve. Furthermore, thickness is a key geometric parameter with clear physical significance and easy control in artificial valve design. Therefore, based on the geometric model of the leaflet of the trilobed artificial biological aortic valve, this application uses the radius of the free edge radius of the leaflet and the leaflet thickness as variable geometric parameters for optimization design.

[0129] In step S110, based on the leaflet geometric model, multiple different leaflet free edge fillet radii are assigned to generate multiple first candidate leaflet models. The leaflet free edge fillet radius is as follows: Figure 3 The radius shown is r, and the obtained first candidate leaflet model is represented in the form of the first candidate leaflet geometric model.

[0130] In one embodiment, multiple first candidate leaflet models have the same leaflet thickness and the same leaflet geometric topology. The leaflet geometric topology refers to the combination of the number of leaflets, fixed edges, ventral surfaces, free edges, and the connections between these parts in the leaflet geometric model. Multiple candidate leaflet models having the same leaflet geometric topology means that the multiple candidate leaflet models maintain consistency in the number of leaflets, the position of fixed edges, the connection between the ventral surfaces and free edges, and the circumferential arrangement of the trefoil, while allowing variations in the free edge radius or leaflet thickness. Therefore, the differences between the multiple candidate leaflet models can be mainly attributed to the assigned variable geometric parameters, facilitating the evaluation of the impact of these variable geometric parameters on leaflet performance.

[0131] In one embodiment, after the parameterized leaflet geometric model is established, the leaflet free edge fillet radius is used as the first geometric design variable, and the parameter scan of the leaflet free edge fillet radius is performed to generate multiple first candidate leaflet models with different free edge geometric shapes.

[0132] Specifically, the multiple different leaflet free edge fillet radii are selected from multiple values ​​within the range of 1 mm to 3.5 mm. Preferably, the leaflet free edge fillet radii include at least three of 1 mm, 1.5 mm, 2 mm, 2.5 mm, 3 mm, and 3.5 mm; more preferably, six first candidate leaflet models are established using the above six fillet radii respectively, to form candidate model groups corresponding to different free edge fillet radii.

[0133] In this embodiment, during the parametric modeling process, all modeling parameters except for the radius of the leaflet's free edge fillet are kept consistent. This includes the leaflet's geometric topology, the generation method of the ventral surface, the leaflet thickness, material parameters, and boundary conditions. Only the radius of the arc forming the free edge is changed to generate multiple first candidate leaflet models. Therefore, each first candidate leaflet model has a unified geometric and material basis, and its simulation evaluation results can primarily reflect the impact of changes in the radius of the leaflet's free edge fillet.

[0134] Preferably, the multiple fillet radii are set in ascending order, that is, multiple first candidate leaflet models with free edge fillet radii of 1 mm, 1.5 mm, 2 mm, 2.5 mm, 3 mm, and 3.5 mm are established sequentially. By setting the above discrete parameters, multiple representative candidate models can be formed within a predetermined range to cover the change process of the free edge from a smaller radius of curvature to a larger radius of curvature, facilitating the analysis of the influence of the free edge fillet radius change on the overall valve performance.

[0135] In some implementations, the number of candidate models can be increased or the interval between adjacent fillet radii can be reduced as needed to further improve the resolution of parameter optimization. However, it is preferable that all candidate fillet radii are within the range of 1 mm to 3.5 mm.

[0136] Please see Figure 7 The figure shows a top view of six sets of tricuspid artificial aortic valves with different leaflet free edge fillet radii in one embodiment of this application. As shown, the six sets of leaflet free edge fillet radii are 1 mm, 1.5 mm, 2 mm, 2.5 mm, 3 mm, and 3.5 mm, respectively. Based on the leaflet geometric model, six first candidate leaflet models are generated for the above six different leaflet free edge fillet radii, respectively referred to as first candidate leaflet model A (labeled as Type 1 in the figure), first candidate leaflet model B (labeled as Type 2 in the figure), first candidate leaflet model C (labeled as Type 3 in the figure), first candidate leaflet model D (labeled as Type 4 in the figure), first candidate leaflet model E (labeled as Type 5 in the figure), and first candidate leaflet model F (labeled as Type 6 in the figure). At this time, the first candidate leaflet model is manifested in the form of the first candidate leaflet geometric model. It should be noted that the values ​​and number of leaflet free edge fillet radii in this application are only examples, and this application is not limited thereto.

[0137] Specifically, after establishing the leaflet geometric model and obtaining the first candidate leaflet geometric model, the material parameters obtained from the uniaxial tensile test of bovine pericardium material are assigned to the first candidate leaflet geometric model, and a hyperelastic constitutive model is used to describe the nonlinear stress-strain relationship of the leaflet material. Since bovine pericardium is a typical biological soft tissue material, it undergoes significant elastic deformation under transvalvular pressure differential during the cardiac cycle, exhibiting obvious nonlinear characteristics in its mechanical response. Therefore, compared to using a linear elastic constitutive model, the hyperelastic constitutive model can more accurately describe the deformation behavior of the leaflet material under large strain conditions, improving the consistency between the structural mechanics analysis results and the actual working conditions.

[0138] In one example, the hyperelastic constitutive model is a second-order Ogden constitutive model. Specifically, based on the stress-strain data obtained from uniaxial tensile tests of bovine pericardium material, the material parameters in the second-order Ogden constitutive model are fitted, and the fitted model parameters are used as constitutive parameters of the leaflet material and input into the finite element analysis model to establish a nonlinear mechanical model of the leaflet material.

[0139] In one specific embodiment, based on the stress-strain data obtained from the uniaxial tensile test, the parameters of the valve material are calibrated in Abaqus using the bivariate Ogden hyperelastic model, as shown in formula (3). The constitutive parameters obtained are shown in Table 1.

[0140] Formula (3)

[0141] In the formula, μ i and α i Let λ1, λ2, and λ3 be the three principal elongation ratios, and i be the term number. In this embodiment, a second-order Ogden constitutive model is used, with i taking values ​​of 1 and 2. i D1 and D2 are compressibility parameters of the material. In this embodiment, both D1 and D2 are 0 to characterize the assumption that the leaflet material is incompressible.

[0142] Table 1. Hyperelastic material parameters used in the second-order Ogden model for valve modeling.

[0143] Because the valve's movement within the body involves highly nonlinear mechanical behavior, including material nonlinearity, large geometric deformation, and contact between leaflets, an 8-node linear reduced integral hexahedral element C3D8R is used to mesh the valve in the finite element simulation to ensure both accuracy and computational efficiency. This element effectively reduces shear locking issues during valve bending deformation, reduces element distortion, and improves computational stability and accuracy. To determine the optimal mesh density, in one embodiment, mesh sensitivity analysis is performed using the aforementioned first candidate leaflet model F. Three meshing schemes are used: mesh sizes of 0.3 mm × 3, 0.2 mm × 3, and 0.2 mm × 4, corresponding to 12042, 20070, and 37056 element numbers, respectively. "0.3 mm × 3" indicates a leaflet feature size of 0.3 mm and 3 layers of elements in the thickness direction; "0.2 mm × 3" and "0.2 mm × 4" indicate a leaflet feature size of 0.2 mm and 3 and 4 layers of elements in the thickness direction, respectively.

[0144] Please see Figure 8 The figure shows the stress distribution cloud map of the valve in the fully open and fully closed states under different grid densities in one embodiment of this application. As shown in the figure, (I) shows the stress distribution of the valve when it is open, and (II) shows the stress distribution of the valve when it is closed, where (a) corresponds to a grid size of 0.3 mm × 3, (b) corresponds to a grid size of 0.2 mm × 3, and (c) corresponds to a grid size of 0.2 mm × 4. The results show that the stress distribution characteristics under the three grid densities are basically the same, and the difference in the maximum stress value is less than 5%, indicating that the grid density has no significant impact on the calculation accuracy. Therefore, in order to effectively balance the calculation cost and the accuracy of the results, the 0.3 mm × 3 grid division scheme was finally selected, and the number of elements of each valve is shown in Table 2.

[0145] Table 2 Number of valve units in the simulation

[0146] In step S120, a multi-index simulation evaluation is performed on multiple first candidate leaflet models. The multi-index simulation evaluation includes at least two of the following evaluations: opening performance evaluation, structural mechanics evaluation based on the dynamic opening and closing process of the leaflet under transvalvular pressure difference, and hemodynamic evaluation based on the aortic flow field model.

[0147] The first candidate leaflet model for multi-index simulation evaluation is represented as a first candidate leaflet finite element model. Therefore, before performing multi-index simulation evaluation on multiple first candidate leaflet models, the process includes constructing a first candidate leaflet finite element model based on each first candidate leaflet geometric model. Specifically, each first candidate leaflet geometric model is assigned the same material properties, a finite element mesh is generated, and contact relationships, analysis steps, inter-leaf pressure differential loads, and boundary constraints are set to obtain the first candidate leaflet finite element model. Thus, each first candidate leaflet finite element model uses the same material parameters, finite element modeling rules, load conditions, and boundary conditions, except for the different leaflet free edge fillet radii, so that the differences in finite element analysis results are mainly attributed to variations in the leaflet free edge fillet radius.

[0148] In one embodiment, the multi-index simulation evaluation includes the orifice performance evaluation, the structural mechanics evaluation, and the hemodynamic evaluation. In other embodiments, the multi-index simulation evaluation may also include any two of the above three types of evaluation. For example, when the leaflet material and aortic flow field conditions have been predetermined, and the main focus is on the leaflet opening capacity and structural load-bearing state, the multi-index simulation evaluation may include orifice performance evaluation and structural mechanics evaluation; when the safety conditions of the leaflet material structure have been predetermined, and the main focus is on the orifice formation capacity and aortic flow field state, the multi-index simulation evaluation may include orifice performance evaluation and hemodynamic evaluation; when all candidate leaflet models meet the preset orifice conditions, the multi-index simulation evaluation may include structural mechanics evaluation and hemodynamic evaluation. For multiple candidate leaflet models at the same stage, the same evaluation type, simulation conditions, and parameter determination rules are used for comparison.

[0149] In one example, all three evaluation types are applied to the same candidate leaflet model. The slit performance evaluation and structural mechanics evaluation are based on the finite element dynamic opening and closing analysis results of the candidate leaflet model; the hemodynamic evaluation is based on computational fluid dynamics analysis of the maximum opening configuration obtained from the finite element dynamic opening and closing analysis and the aortic flow field model. Therefore, for the same candidate leaflet model, corresponding evaluations can be performed from three aspects: opening function, structural mechanics response, and aortic flow field response. Specifically, the slit performance evaluation reflects the opening and closing function of the artificial valve, the structural mechanics evaluation reflects the mechanical reliability and durability of the leaflet tissue, and the hemodynamic evaluation reflects the impact of the artificial valve on the aortic blood flow environment. By comprehensively considering the results of these three evaluation types, valve function, structural safety, and blood flow quality can be simultaneously taken into account, avoiding the performance degradation caused by optimizing only a single performance indicator, thus obtaining an artificial aortic valve design with superior overall performance.

[0150] Among them, orifice performance evaluation refers to the simulation evaluation process used to assess the ability of a candidate valve leaflet model to form a blood flow channel in the open state. Orifice performance evaluation may include determining the orifice area formed when the candidate valve leaflet model reaches its maximum opening degree during the cardiac cycle. Orifice performance evaluation is used to reflect the impact of changes in valve leaflet geometric parameters on valve opening adequacy and blood flow channel formation ability.

[0151] During a cardiac cycle, as the pressure gradient increases, the valve slowly opens, reaching its maximum opening and then gradually closing as the pressure gradient decreases. Excessive displacement or strain may indicate that the material is too soft or excessively deformed, resulting in uneven opening or insufficient effective opening area, affecting blood flow. Therefore, it is necessary to assess the strain and displacement distribution of the valve under transvalvular pressure gradients. Please refer to [link to relevant documentation]. Figure 9The figure illustrates the dynamic changes in valve displacement during a cardiac cycle in one embodiment of this application. As shown, during the opening phase, the increased transvalvular pressure gradient causes the edges of adjacent leaflets to gradually separate, and the ventral side of the leaflet slowly bulges upward, reaching its maximum displacement at time t2. Subsequently, the displacement value gradually decreases from the free end to the fixed edge. During the closing phase, the valve begins to retract from the central region, and the opening area gradually decreases until it is completely closed. The maximum deformation of each model is located in the middle of the free edge, indicating that this region is the most significant part of the valve's dynamic response. Different fillet radius models exhibit similar opening and closing patterns throughout the cycle, but there are differences in local morphological changes. As the fillet radius increases, the transition of the valve connection area and the free edge becomes smoother, the valve orifice contour becomes more regular during the opening phase, and the free edge closure becomes more continuous during the closing phase. Especially at times t2 and t3, the larger fillet radius model exhibits better motion coordination and contour continuity, indicating that fillet optimization can improve the local geometric discontinuities in the valve's dynamic deformation to a certain extent. By the end of the closure phase, each model forms a relatively complete tricuspid closure structure, but the larger fillet radius model has a relatively more stable closure morphology. Overall, increasing the radius of the rounded corner helps improve the morphological smoothness and motion consistency during the dynamic evolution of the valve, but its improvement effect tends to weaken within a larger radius.

[0152] In one embodiment, the apical performance evaluation includes determining the geometric opening area of ​​the first candidate leaflet model in the open state; the geometric opening area is the valve orifice opening area formed when the first candidate leaflet model reaches its maximum opening degree during the cardiac cycle.

[0153] The geometric orifice area (GOA) refers to the area of ​​the valve orifice region formed by the free edges of the three leaflets when the candidate leaflet model reaches its maximum opening degree during the cardiac cycle. GOA can be used to evaluate the ability of the leaflet model to allow blood flow when it is open. In some embodiments, the valve orifice morphology can be captured at the moment the candidate leaflet model reaches its maximum opening degree, and the geometric orifice area can be determined based on the valve orifice boundary.

[0154] As mentioned above, a larger GOA value indicates a more ideal degree of valve opening during a complete cardiac cycle. Under otherwise identical conditions, the geometric opening area affects the blood flow channel area and blood flow capacity of the valve orifice; however, the actual blood flow is also influenced by factors such as transvalvular pressure gradient, fluid properties, and flow resistance. At time t2, the circular opening area formed between the valves reaches its maximum value; therefore, it is necessary to calculate the geometric opening area of ​​different models at this time. Please refer to [link to relevant documentation]. Figure 10The diagram shows a schematic of the calculation of the valve geometric opening area in one embodiment of this application. Taking the valve fully open as an example: First, the moment when the opening area is the largest is determined, and the shape of the valve opening is captured; then, the screenshot is imported into ImageJ software, and the suture edge (DAA) and free edge (Lf) of the valve are obtained through boundary extraction; finally, the opening area is calculated.

[0155] Using the above method, the geometric opening areas of candidate leaflet models A, B, C, D, E, and F were calculated. It was found that the geometric opening area increases with the increase of the corner radius. Candidate leaflet model F has the largest opening area. This indicates that valves with larger corner radii have larger geometric opening areas.

[0156] Structural mechanics evaluation refers to the process of evaluating the structural response of a candidate leaflet model during the dynamic opening and closing process under the action of transvalvular pressure differential, based on finite element analysis or other structural mechanics simulation methods. In one embodiment, during structural mechanics evaluation, finite element analysis is first performed on the leaflet geometric model or the candidate leaflet model to obtain a finite element model, and then structural mechanics evaluation is performed based on the finite element model. Structural mechanics evaluation may include determining one or more of the following parameters of the candidate leaflet model during the cardiac cycle: maximum principal stress, maximum principal strain, von Mises equivalent stress (also referred to as equivalent stress in this text), and maximum shear stress. Structural mechanics evaluation is used to characterize the influence of changes in leaflet geometric parameters on stress concentration, local deformation, overall stress distribution, and structural stability. Transvalvular pressure differential refers to the difference in pressure between the two sides of the leaflet. For a tricuspid aortic valve, the transvalvular pressure differential can be the difference between the ventricular side pressure and the aortic side pressure. The transvalvular pressure differential drives the leaflets to open and close during the cardiac cycle.

[0157] In one embodiment, the structural mechanics evaluation includes: applying a transvalvular pressure differential load during the cardiac cycle to the first candidate leaflet model, and performing finite element dynamic opening and closing analysis on the first candidate leaflet model.

[0158] In one example, the transvalvular pressure differential load is a pressure load applied to a candidate leaflet model based on the transvalvular pressure differential. This transvalvular pressure differential load can be a time-history load that varies with the cardiac cycle, used to simulate the dynamic opening and closing process of the leaflet within one cardiac cycle. In finite element dynamic opening and closing analysis, the transvalvular pressure differential load can be applied to the leaflet surface, and the displacement, stress, strain, and deformation results of the leaflet can be obtained by combining fixed edge constraints. The transvalvular pressure differential load can be constructed based on the normal physiological pressure curve of the human body, or it can be personalized based on the clinical test data of the target patient, enabling the finite element analysis process to more realistically reflect the stress situation of the artificial aortic valve under actual working conditions.

[0159] In one example, the finite element dynamic opening and closing analysis refers to the simulation analysis process of using the finite element method to simulate the opening and closing process of a candidate leaflet model during the cardiac cycle under transvalvular pressure differential load, and to obtain the mechanical properties of the leaflet structure. Finite element dynamic opening and closing analysis can be used to obtain the stress, strain, displacement, and deformation distribution of the candidate leaflet model at multiple moments or throughout the entire cardiac cycle.

[0160] In one example, the finite element dynamic opening and closing analysis can be used to obtain the displacement field, velocity field, stress field, strain field, contact pressure distribution, strain energy density and reaction force of the first candidate leaflet model at each moment during the cardiac cycle, and to evaluate the uniformity of force, stress concentration, deformation coordination and potential fatigue risk of the leaflet during the opening and closing process.

[0161] In engineering practice, four strength theories are used to determine whether a material will fail under complex stress conditions. Among them, the maximum principal stress, maximum principal strain, maximum shear stress, and equivalent stress are directly related to material failure. Therefore, in one embodiment, the structural mechanics evaluation includes determining the maximum principal stress, maximum principal strain, von Mises equivalent stress, and maximum shear stress of the first candidate leaflet model during dynamic opening and closing. In one example, the structural mechanics evaluation includes obtaining the time-history peak values ​​and / or spatial average values ​​of the maximum principal stress, maximum principal strain, von Mises equivalent stress, and maximum shear stress during the cardiac cycle.

[0162] In this application, the spatial average value within the leaflet region refers to the value obtained by spatially averaging the structural mechanical field variables corresponding to each finite element element or sampling point within the three leaflet regions of the candidate leaflet model at each calculation time of the cardiac cycle. The spatial average values ​​at each calculation time constitute the spatial average time history curve of the corresponding index, and the peak value can be further determined from the spatial average time history curve. Therefore, in this paper, "average maximum principal stress," "average maximum principal strain," "average maximum equivalent stress," and "average maximum shear stress" respectively represent the spatial average results of the corresponding field variables within the leaflet region. For details, please refer to... Figure 11a and Figure 11b , Figure 11a The graph shown is a curve illustrating the change of the maximum principal stress on the valve over time in one embodiment of this application. Figure 11b The figure shows the curve of the average maximum principal stress on the valve over time in one embodiment of this application. As shown, the maximum principal stress and average maximum principal stress of different fillet radius models exhibit periodic changes, with stress peaks mainly occurring during the on / off transition phase and the later loading phase. The fillet radius variation has a limited impact on the overall evolution of the valve principal stress, but it has a certain moderating effect on transient peaks and local fluctuations. Models with smaller fillet radii show higher maximum principal stress peaks and more pronounced fluctuations in the early stage of the cycle, indicating that sharper geometric transitions are more prone to local stress concentration. As the fillet radius increases, the stress curve gradually flattens out. The first candidate leaflet model F shows a relatively low stress level in the later main peak stage, indicating that a larger fillet radius helps improve the continuity of stress in the connection area and reduce local high stress concentration. In contrast, the differences in average maximum principal stress among the models are small, indicating that fillet optimization has a limited impact on the overall average stress level, and its effect is mainly reflected in the local stress regulation.

[0163] Please see Figure 12The figure shows the maximum principal stress cloud map of the first candidate leaflet model at different times in one embodiment of this application. As shown in the figure, the distribution of high stress areas is generally consistent among the models during dynamic opening and closing, mainly located at the free edge, attachment edge, junction area, and closure anastomosis area of ​​the valve, but the degree of local concentration varies. At times t1 and t2, the first candidate leaflet models A to C show more significant stress concentration at the valve edge and connecting transition area, while the first candidate leaflet models D and E show some relief, and the first candidate leaflet model F has the most gradual distribution. At time t3, the principal stress of each model decreases overall, and the difference between groups weakens, but the first candidate leaflet models D and F are closer to a uniform low stress state. By time t4, the high principal stress is concentrated again in the anastomosis area and connecting transition area, with the first candidate leaflet models A to C being the most significant, and the first candidate leaflet model F being the weakest. Overall, it shows that as the radius of curvature increases, the local principal tensile stress concentration gradually decreases, and the first candidate leaflet model F performs the best.

[0164] As shown in Table 3, as the fillet radius increases from candidate leaflet model A to candidate leaflet model F, both the peak value of the maximum principal stress and the average maximum principal stress of the valve decrease, indicating that increasing the fillet radius helps improve the stress state of the valve, especially reducing local principal tensile stress concentration. Regarding the peak value of the maximum principal stress for each model, candidate leaflet model F decreases by approximately 28.0% compared to candidate leaflet model A; regarding the average maximum principal stress, candidate leaflet model F decreases by approximately 11.7% compared to candidate leaflet model A. These results indicate that increasing the fillet radius not only reduces local peak stress but also improves the overall stress uniformity of the valve to a certain extent.

[0165] Table 3 Maximum principal stress of the aortic valve

[0166] Please see Figure 13a and 13b , Figure 13a The graph shown is a curve illustrating the maximum principal strain experienced by the valve over time in one embodiment of this application. Figure 13bThe figure shows a curve illustrating the variation of the average maximum principal strain over time in one embodiment of this application. As shown, the maximum principal strain and average maximum principal strain of different fillet radius models exhibit periodic variations. In contrast, candidate leaflet models A to C show higher principal strain levels during the pre-cardiac cycle and peak phase, indicating that leaflets with smaller fillet radii are more prone to local tensile deformation. The principal strain levels of candidate leaflet models D and E are reduced, while candidate leaflet model F shows the lowest principal strain for most periods, indicating that a larger fillet radius can improve the geometric continuity of the leaflet edge and connecting transition region, reducing local excessive elongation during dynamic opening and closing.

[0167] Please see Figure 14 The figure shows the maximum principal strain cloud map of the first candidate leaflet model at different times in one embodiment of this application. As shown, the principal strain of each model is mainly concentrated in the free edge, attachment edge, and transition region of the valve, and the degree of local concentration varies significantly under different fillet radii. At times t1 and t2, the high principal strain regions of the first candidate leaflet models A to C are more prominent, the first candidate leaflet models D and E are in a transitional state, and the first candidate leaflet model F has the most gradual distribution. At time t3, the overall principal strain of each model is relatively low, and the differences between groups are weakened, but the first candidate leaflet models D to F are closer to a uniform low strain state. By time t4, the first candidate leaflet models A to C still have obvious principal strain concentration in the mating zone and transition region, while the first candidate leaflet model F is the weakest. Overall, as the fillet radius increases, the local excessive elongation gradually decreases, and the first candidate leaflet model F performs best in improving deformation compatibility.

[0168] Table 4 shows the peak and average maximum principal strain experienced by the valve during one cardiac cycle. For the peak maximum principal strain of each model, compared to the first candidate leaflet model A, the peak maximum principal strain of the first candidate leaflet model F decreased by approximately 18.8%. For the peak average maximum principal strain of each model, compared to the first candidate leaflet model A, the peak average principal strain of the first candidate leaflet model F decreased by approximately 16.7%. Although the average strain value fluctuated slightly between the first candidate leaflet models C and D, it still showed an overall decreasing trend, indicating that the increased corner radius not only reduced local extreme stretching but also lowered the overall principal tensile deformation level of the leaflet.

[0169] Table 4 Maximum principal strain of the aortic valve

[0170] Please see Figure 15a and 15b , Figure 15aThe graph shown is a curve illustrating the change of the maximum equivalent stress on the valve over time in one embodiment of this application. Figure 15b The figure shows the curve of average maximum equivalent stress versus time in one embodiment of this application. As shown, the maximum equivalent stress and average maximum equivalent stress of different rounded corner radius models all change periodically with time, and the overall variation pattern is basically consistent. In contrast, the first candidate leaflet models A to C show relatively high equivalent stress levels in the early and late peak stages of the cardiac cycle, indicating that the valve is more prone to complex stress concentration under smaller rounded corner radii. The stress levels of the first candidate leaflet models D and E are somewhat mitigated, while the first candidate leaflet model F shows the lowest or near-lowest equivalent stress for most periods, indicating that a larger rounded corner radius can improve the geometric continuity of the valve edge and connecting transition area, and reduce the combined stress superposition caused by the combined effects of tension, bending and contact during dynamic circulation. Overall, increasing the rounded corner radius helps to reduce the valve stress level and improve dynamic stress stability, with the first candidate leaflet model F performing the best.

[0171] Please see Figure 16 The figure shows the equivalent stress cloud map of the first candidate leaflet model at different times in one embodiment of this application. As shown, the high stress areas of the first candidate leaflet models A to F are mainly distributed in the leaflet edge, attachment area, connection transition area, and closed engagement area, but there are differences in the stress distribution range and concentration degree among different models. At times t1 and t2, the high equivalent stress areas of the first candidate leaflet models A and B are larger and the stress values ​​are higher, followed by the first candidate leaflet models C and D, indicating that the opening stage is more prone to the formation of compound stress superposition under a small rounded corner radius; while the stress distribution of the first candidate leaflet models E and F is more uniform, with the weakest local high stress concentration in the first candidate leaflet model F. At time t3, the overall equivalent stress of each model is significantly reduced. The first candidate leaflet models A to C still retain a certain degree of stress residue, while the stress distribution of the first candidate leaflet models D to F is more uniform and the overall stress level is lower. At time t4, during the closed loading stage, the higher-order stresses of each model re-concentrated in the mating region and the transitional area. The stress concentration was most pronounced in candidate leaflet models A through C, somewhat mitigated in candidate leaflet models D and E, and most evenly distributed in candidate leaflet model F. Overall, as the corner radius increased, the overall load-bearing strength of the valve gradually improved. Candidate leaflet models A through C experienced the heaviest overall load, while candidate leaflet models D and E were in a transitional state. Candidate leaflet model F showed the best performance in improving stress distribution and reducing peak stress.

[0172] Table 5 shows the peak and average maximum equivalent stress of each model. Regarding the peak maximum equivalent stress of each model, compared to the first candidate leaflet model A, the peak value of the first candidate leaflet model F is reduced by approximately 18.0%. As the fillet radius increases, the peak equivalent stress decreases significantly, indicating that a larger fillet radius has a more pronounced mitigating effect on localized composite stress concentration. Regarding the average maximum equivalent stress of each model, the decreasing trend is more gradual. Compared to the first candidate leaflet model A, the average maximum equivalent stress of the first candidate leaflet model F is reduced by approximately 13.6%. It can be seen that the decrease in average maximum equivalent stress is not as significant as the decrease in peak stress in the first candidate leaflet model F, but overall it shows a continuous improvement, indicating that as the fillet radius increases, the overall composite load state of the leaflet gradually decreases, and the stress distribution tends to be more uniform.

[0173] Table 5 Maximum equivalent stress of the aortic valve

[0174] Please see Figure 17a and 17b , Figure 17a The graph shown is a curve illustrating the change of maximum shear stress on the valve over time in one embodiment of this application. Figure 17b The figure shows the curve of the average maximum shear stress on the valve over time in one embodiment of this application. As shown, the first candidate leaflet models A and B exhibit higher shear stress levels and more obvious fluctuations in the early and late peak stages of the cycle, indicating that the valve edge and connecting transition area are more prone to local slippage and tangential deformation incoordination during rapid opening and closing with a small fillet radius. The first candidate leaflet models C and D are improved compared to the previous two groups, but still have a certain fluctuation plateau in the early and middle stages, indicating that there is still room for optimization in their tangential deformation coordination. In contrast, the first candidate leaflet models E and F have lower overall shear stress, with the first candidate leaflet model F maintaining the lowest level for most periods, indicating that a larger fillet radius helps to smooth the valve edge transition and reduce shear stress concentration caused by local bending, torsion, and contact incoordination. Therefore, as the fillet radius increases, the tangential force disturbance during the dynamic opening and closing of the valve gradually weakens, and the first candidate leaflet model F performs best in improving tangential deformation coordination.

[0175] Figure 18The figure shows the maximum shear stress cloud map of the first candidate leaflet model at different times in one embodiment of this application. As shown, the shear stress of each model is mainly concentrated in the free edge, attachment edge and connection transition region of the valve, but the degree of local concentration varies significantly. At times t1 and t2, the high shear stress region of the first candidate leaflet models A to C is more significant, the first candidate leaflet models D and E are in a transition state, and the first candidate leaflet model F has the most gradual distribution. At time t3, the overall shear stress of each model is lower, and the difference between groups weakens, but the first candidate leaflet models D to F are closer to a uniform low stress state. At time t4, the first candidate leaflet models A to C still show obvious shear stress concentration in the occlusal region and connection transition region, while the first candidate leaflet model F is the weakest. Overall, as the radius of the fillet increases, the tangential force disturbance gradually weakens, and the first candidate leaflet model F performs best in improving the force coordination.

[0176] Table 6 lists the maximum shear stress of the aortic valve during one cardiac cycle. In terms of peak shear stress, compared to the first candidate leaflet model A, the first candidate leaflet model F shows a decrease of approximately 24.7%. In terms of average maximum shear stress, the change is more stable, and it shows an overall decreasing trend. Compared to the first candidate leaflet model A, the first candidate leaflet model F shows a decrease of approximately 10.2%. Although the decrease in average value is not as significant as that of peak value, its trend is more stable, indicating that the increased corner radius not only weakens local extreme shear stress but also improves the overall tangential stress distribution of the leaflet, making the tangential load transfer more uniform during dynamic opening and closing.

[0177] Table 6 Maximum shear stress of the aortic valve

[0178] Since Spearman rank correlation analysis has a wide range of applications and is mainly used to solve problems related to categorical and continuous variables, it does not have special requirements for the distribution type of variables and the sample size. Based on this, in order to better analyze the influence of valve corner radius on mechanical parameters, this application uniformly selects multiple sample points in the valve free edge region and uses Graphpad Prism 9.1.0 to perform t-distribution two-tailed Spearman correlation analysis.

[0179] In one specific implementation, multiple sample points were uniformly selected in the valve free edge region, and the principal stress, von Mises equivalent stress, shear stress, and strain values ​​were collected at each point. 54 sample points were selected for each model, resulting in a sample size of 324. The valve leaflet geometry model was set to six levels according to the radius of the fillet, and Spearman correlation analysis was performed in Graphpad Prism 9.1.0 to evaluate the correlation between the fillet radius level and each mechanical parameter. The results showed that the valve fillet radius was negatively correlated with principal stress, principal strain, shear stress, and equivalent stress, and all correlations reached statistical significance (P < 0.001, N = 324). Among these, the fillet radius showed the strongest correlation with equivalent stress (r = -0.197), with correlation coefficients of -0.191 for both principal strain and shear stress, while the correlation with principal stress was relatively weak (r = -0.111). Overall, although the absolute values ​​of all correlation coefficients are less than 0.2, indicating a weak correlation, their directions are consistent and statistically significant, suggesting that the local stress and strain levels of the valve decrease with increasing valve radius. This result indicates that increasing the radius helps alleviate local stress concentration in the valve, improves the stress distribution, and thus optimizes its mechanical microenvironment.

[0180] Hemodynamic evaluation refers to the process of evaluating the blood flow response after a candidate valve leaflet model is incorporated into the aortic flow field model, based on the aortic flow field model and computational fluid dynamics analysis. Hemodynamic evaluation may include determining one or more of the following: blood flow velocity, wall pressure, and wall shear stress. It is used to characterize the effects of changes in valve leaflet geometry on the valve orifice jet, velocity distribution within the aorta, wall pressure, and tangential wall load.

[0181] In one embodiment, the hemodynamic evaluation includes: reconstructing the aortic flow field model based on aortic medical imaging data, and respectively placing the first candidate leaflet model and the second candidate leaflet model into the aortic flow field model for computational fluid dynamics analysis.

[0182] Computational fluid dynamics (CFD) analysis refers to the process of numerically calculating the flow state of blood in an aortic flow field model based on an aortic flow field model, candidate valve models, and preset fluid boundary conditions. CFD analysis can be used to obtain hemodynamic parameters such as blood flow velocity, wall pressure, and wall shear stress.

[0183] The aortic medical imaging data can be real clinical data, such as CT images of patients with aortic insufficiency. An aortic model is extracted from the CT images and a mesh is generated to provide a physiological three-dimensional model reflecting the hemodynamic parameters at various points on the aortic arch. The aortic flow field model construction process includes: first, extracting the aortic model based on the CT images; then, performing Boolean operations between the obtained aortic model and the aforementioned valve leaflet geometric model to obtain the aortic flow field model.

[0184] Specifically, using CT images of a patient with aortic insufficiency, Mimics 21.0 (Materialise, Belgium), a commonly used software for inverse modeling of clinical cardiovascular diseases, was used to extract the aortic arch. After importing the DICOM format image, the software displayed coronal, sagittal, and cross-sectional views. Alignment between the image and the components of the 3D model was achieved through orientation and position adjustments. For example, the grayscale range was first adjusted to highlight the aorta's location and direction, followed by threshold segmentation. To save time, a region-growing-based segmentation method was used, distinguishing the target from the background based on pixel attributes. The specific steps included: selecting a starting point within the aorta, merging similar feature regions as new starting points, and expanding outwards. This method can extract the aorta relatively completely, but it easily produces redundant structures. Based on this, the heart and redundant branch arteries were trimmed as needed, as shown in the figure. For any internal voids that might be generated by region growing, Mimics' "Smart fill" was used to fill and smooth the area, resulting in a complete and smooth aortic model. Geomagic was then used for trimming, smoothing, and filling, and finally Spaceclaim was used to process the fluid model. The inlet is located in the left ventricular outflow tract and is the channel through which blood flows into the aorta; outlet 1 is set as the outlet of the descending thoracic aorta; outlets 2 to 4 correspond to the outlets of the brachiocephalic trunk, the left common carotid artery, and the left subclavian artery, respectively.

[0185] Please see Figure 19 The figure shows a geometric model of the aortic valve leaflets at peak systolic time in one embodiment of this application. As shown, the geometric changes of the heart valve at the peak systolic time t2 within a cardiac cycle produce differences in deformation characteristics. This is the most important typical moment in aortic CFD analysis. Therefore, using the fully open valve at the peak systolic time t2 as a reference, and based on its size and opening morphology, a three-dimensional reconstruction is performed in Abaqus to obtain the following... Figure 19The valve leaflet geometry model is shown. It should be noted that in this embodiment, the structural mechanics evaluation and hemodynamic evaluation are performed sequentially using simulation, rather than in a two-way fluid-structure interaction calculation within the same solution process. Specifically, firstly, the maximum opening configuration of each candidate valve leaflet model at the peak systolic time is obtained through finite element dynamic opening and closing analysis; subsequently, this maximum opening configuration is converted or reconstructed into a static valve leaflet geometry model for computational fluid dynamics analysis, and this static valve leaflet geometry model is placed within the same aortic fluid domain. Each candidate valve leaflet model uses the same aortic geometry, fluid properties, and inlet and outlet boundary conditions, so that the hemodynamic evaluation results mainly reflect the effects of changes in the geometric parameters of the candidate valve leaflet models.

[0186] The reconstructed aortic model is presented in solid form to characterize the blood flow within the aorta. However, since the leaflet structure is located inside the aorta, a Boolean operation is performed between the fluid domain corresponding to the aortic lumen and the solid leaflet model. The volume occupied by the solid leaflet is subtracted from the fluid domain to form a geometric model of the aortic flow field with leaflet boundaries and valve orifice passages.

[0187] Before performing CFD simulations, the model must first be meshed, i.e., the flow field must be discretized. The meshing method and the number of mesh points directly affect the accuracy and reliability of the CFD calculation results. To obtain a high-quality finite element volume mesh, the 3D model file created above is imported into ANSYS 2022 Fluent Meshing for meshing. The basic steps include: importing the geometric model and checking the surfaces; defining the inlet, outlet, and walls of the model and setting boundary conditions; meshing the surfaces; and meshing the volume.

[0188] Mesh quality determines computational convergence and accuracy. To determine suitable mesh size and number, a mesh independence test was performed. Using the maximum wall shear stress (WSSmax) of the first candidate leaf model A as the standard, different element sizes were used to generate different numbers of meshes. A mesh was considered feasible when the relative error of WSSmax was < 5%. The mesh independence test results showed that when the mesh number reached approximately 849,200, the relative error was 3.59% (< 5%), and the residual converged to 1 × 10⁻⁶. -4At that time, the inlet and outlet flow pressures were stable. Therefore, all models used a uniform mesh: size 1 mm, wall boundary layer (first layer 0.02 mm, 3 layers, increasing by 1.2). Under this scheme, the WSS results all converged, and the standard mesh parameters were defined. Wall shear stress (WSS) refers to the tangential stress acting on the wall or a pre-defined wall region of the aortic flow field model during blood flow. WSS can be used to evaluate the scouring effect of blood flow on the aortic wall, tangential load, and near-wall flow state.

[0189] In one example, the aortic flow field model includes a left ventricular outflow tract inlet and multiple aortic outlets, and the computational fluid dynamics analysis includes using the left ventricular outflow tract inlet as a velocity inlet and the multiple aortic outlets as pressure outlets.

[0190] In one specific implementation, one of the commonly used boundary conditions in large vessel CFD simulations is the velocity-pressure boundary. The flow rate at the peak systolic velocity was set to 385.5 g / s, equivalent to a blood flow velocity of approximately 1.1 m / s, and this was used as the inlet velocity boundary condition for the model. The pressure change waveforms at the four outlets of the aorta were simulated using a lumped parameter model, and the pressure values ​​at time t2 are shown in Table 7.

[0191] Table 7 Boundary Conditions

[0192] In addition, to solve fluid models using CFD methods, besides boundary conditions, specific initial conditions must also be defined to obtain an accurate solution for the model. Furthermore, it is necessary to consider whether the blood flow is laminar or turbulent. The dimensionless parameter—the Reynolds number (Re)—is a similarity criterion number in fluid mechanics that characterizes the influence of viscosity. The blood flow state is often distinguished by Re, which is the ratio of inertial force to viscous force, and its definition is shown in formula (4):

[0193] Formula (4)

[0194] Where ρ is the blood density (kg / m³) 3 d is the blood vessel diameter (m), μ is the hemodynamic viscosity (kg / (m·s)), and v is the blood velocity (m / s). Blood flow is influenced by both inertial and viscous forces. When Re is low, viscous forces dominate, disturbances are reduced, resulting in laminar flow; when Re > 2000, inertial forces dominate, disturbances are enhanced, resulting in turbulent flow. Blood is viscous, exhibiting significant viscosity at low shear rates and constant viscosity at high shear rates, thus it is considered a Newtonian fluid. The bulk modulus of elasticity of blood and water are similar, both being incompressible fluids.

[0195] In one example, in the computational fluid dynamics analysis, blood is set as a homogeneous, incompressible, isotropic Newtonian fluid, and the aortic wall is set as a rigid, non-slip wall. Specifically, for large blood vessels like the aorta, in the steady flow simulation of this embodiment, blood is set as a homogeneous, incompressible, isotropic Newtonian fluid with a density of 1050 kg / m³. 3 The dynamic viscosity is 0.0035 Pa·s, the inlet velocity is 1.1 m / s, and the Poisson's ratio is 0.45. The Reynolds number calculated according to formula (4) is greater than 2300, so a turbulence model is used for solution, and the Realizable k-ε model is selected. The aortic vessel wall is set as a rigid wall, and no-slip boundary conditions are used.

[0196] The core of hemodynamics is solving the equations of mass conservation and momentum conservation, namely the continuity equation and the Navier-Stokes equation.

[0197] Mass conservation: For incompressible fluids (density ρ is constant), under steady flow conditions, the flow rate is conserved along the pipe.

[0198] Formula (5)

[0199] In the formula, Q is the volumetric flow rate, A is the cross-sectional area, and V is the cross-sectional average velocity.

[0200] Conservation of momentum: In a rectangular coordinate system, the Navier-Stokes equations for incompressible fluids are:

[0201]

[0202]

[0203] Formula (6)

[0204] In the formula, u, v, and w represent the blood velocity components (m / s) in the x, y, and z directions, respectively, and ρ is the blood density (kg / m³). 3 μ is the hemodynamic viscosity (Pa·s), and p is the vessel wall pressure (Pa).

[0205] In this application, numerical simulations were conducted using ANSYS 2022 R1 Fluent software, and the SIMPLE (semi-implicit) algorithm was selected for solution. The convection terms in the momentum and continuity equations were discretized using a second-order upwind scheme. During the calculations, the convergence criterion for all residual values ​​was set to 1×10⁻⁶. -4 When the residual is below this value and the monitored physical quantity no longer changes significantly, the simulation result is considered to have converged.

[0206] Based on the above aortic flow field geometric model, in one embodiment, the hemodynamic evaluation includes determining the blood flow velocity, wall pressure, and wall shear stress corresponding to the first candidate leaflet model after it is placed into the aortic flow field model.

[0207] Please see Figure 20 The figure shows a typical planar position of the aorta in one embodiment of this application. As shown in the figure, A is the proximal ascending aorta, B is the middle of the ascending aorta, C is the distal ascending aorta, D is the proximal aortic arch, E is the middle of the aortic arch, F is the distal aortic arch, S1 represents segment OA; S2 represents segment AB; S3 represents segment BC; S4 represents segment CD; and S5 represents segment DE.

[0208] Please see Figure 21 The figure shows the blood flow pattern cloud map of valves with different fillet radii during the peak systolic phase in the same aortic flow field in one embodiment of this application. As shown in the figure, the upper part is the front view and the lower part is the rear view. The change in fillet radius mainly changes the formation mode and spatial distribution of the jet near the valve orifice. During the peak systolic phase, the overall flow pattern of each model is basically consistent, with high velocity values ​​mainly located at the valve orifice outlet and the local tortuous transition region. As the fillet radius increases, the jet near the valve orifice gradually tends to be more concentrated and straight, the local high-velocity band attached to the wall and the streamline deflection are weakened, and the uniformity of velocity distribution is improved. Among them, the first candidate leaflet model E and the first candidate leaflet model F show more stable central jet characteristics.

[0209] Please see Figure 22 The figure shows a schematic diagram of the maximum blood flow velocity in the same aortic flow field during the peak systolic period of valves with different fillet radii in one embodiment of this application. As shown in the figure, the Vmax of the first candidate leaflet model A and the first candidate leaflet model B is 1.44 m / s. When the fillet radius increases to 2.0 and 2.5 mm, Vmax decreases to 1.39 m / s and 1.38 m / s, respectively. After further increasing to 3.0 and 3.5 mm, it rises back to 1.48 m / s and 1.46 m / s, with an overall fluctuation range of 1.38–1.48 m / s, a relative change of about 7%. It can be seen that the influence of the fillet radius on the maximum blood flow velocity is relatively limited and exhibits a non-monotonic change.

[0210] To quantify blood flow velocity changes in different models, the maximum and average flow velocities at typical locations on the aortic wall during peak systole were measured. Please refer to [link / reference]. Figure 23a and 23b , Figure 23a This diagram illustrates the maximum blood flow velocity in a segmented aortic region according to one embodiment of this application. Figure 23bThe figure shows a schematic diagram of the average blood flow velocity in the aortic region in one embodiment of this application. As shown, although the first candidate leaflet models A to F have similar overall flow trends, they still have certain differences in maximum blood flow velocity. The first candidate leaflet model D consistently exhibits a high velocity level, indicating that this type is more likely to form local high-speed jets and a stronger flow impact effect; the first candidate leaflet models C and B are next, and their velocity levels are relatively close to those of the first candidate leaflet model D; the first candidate leaflet model A is in an intermediate state; while the first candidate leaflet models E and F exhibit lower velocity levels, with the first candidate leaflet model F having the lowest, indicating that its local flow is relatively gentle and is less likely to cause a sharp increase in WSS.

[0211] Please see Figure 24 The figure shows cross-sectional cloud maps of blood flow velocity in the same aorta for heart valves with different fillet radii during peak systole in one embodiment of this application. As shown, five cross-sections of the aortic lumen were created from the proximal end of the ascending aorta (section A) to the middle end of the aortic arch (section E), indicating that there are significant differences in the cross-sectional distribution of blood flow velocity during peak systole for different fillet radius models. The first candidate leaflet model D exhibits the most significant eccentric high-velocity jet, followed by the first candidate leaflet models B and C, with the first candidate leaflet model A in the middle, while the first candidate leaflet models E and F have smaller high-velocity jet ranges. Among them, the first candidate leaflet model F has a more uniform velocity distribution in each cross-section, indicating that it has the least disturbance to downstream flow and a more stable overall hemodynamic performance.

[0212] Wall pressure is an important indicator for evaluating aortic hemodynamic load. Its distribution characteristics are closely related to valve geometry; high pressure concentration often indicates an increased risk of vascular wall remodeling and related complications. Please refer to [link to relevant documentation]. Figure 25 The figure shows the vascular pressure distribution cloud map at the peak systolic time for valves with different fillet radii in one embodiment of this application. As shown, under the condition of smaller fillet radii (first candidate leaflet model A, first candidate leaflet model B), the wall pressure is more concentrated, especially on the inner side of the valve and in the high curvature region. As the fillet radius increases (first candidate leaflet models C to F), the wall pressure distribution gradually becomes more uniform, and the local high pressure area is significantly reduced. Especially in first candidate leaflet models E and F, the pressure peak is further reduced, and the pressure concentration phenomenon in the curvature region is significantly alleviated. Therefore, it is shown that a larger fillet radius helps to improve the pressure distribution on the valve surface, reduce local abnormal load, and thus improve the hemodynamic stability of the valve.

[0213] Please see Figure 26The figure shows a schematic diagram of the maximum wall pressure at the peak systolic time for valves with different fillet radii in one embodiment of this application. As shown in the figure, the effect of different fillet radii on the maximum pressure of the aortic wall is relatively stable. The variation range of the maximum pressure values ​​for all types is very small, all within 1.29 × 10⁻⁶. 4 Pa up to 1.31 × 10 4 The pressure levels were relatively stable, with the radius of curvature between 100 Pa and 100 Pa. This stability reflects the relatively weak impact of changes in the fillet radius on the overall fluid pressure, with minimal fluctuations in the maximum pressure. However, even under these stable pressure conditions, different fillet radii exhibited different pressure distribution characteristics. Smaller fillet radii, due to higher local pressure peaks, could lead to local flow instability and potentially increased hemodynamic load on the local aortic wall. As the fillet radius increased, the pressure distribution became smoother, and the area of ​​local pressure concentration decreased, indicating that a larger fillet radius is beneficial for improving the wall pressure distribution in the aortic flow field. Furthermore, as the fillet radius increased, the flow became more stable, and the pressure concentration area decreased, suggesting that a larger fillet radius helps improve valve durability and reduce damage to the valve material from local pressure.

[0214] Please see Figure 27 The figure shows a schematic diagram of the average wall pressure of each model in the aortic segment region in one embodiment of this application. As shown in the figure, the average wall pressure of each model changes along the flow path in a basically consistent manner, with small differences between the types, indicating that the change in the radius of curvature has a limited impact on the overall wall pressure level. In contrast, although the first candidate leaflet model F did not significantly reduce the peak wall pressure, its pressure distribution was smoother and the fluctuations along the flow path were smaller, reflecting better flow field uniformity and lower downstream disturbance.

[0215] Wall shear stress (WSS) is the tangential frictional force exerted by blood flow on the vessel wall. Its magnitude is directly related to the blood flow velocity gradient. High WSS distributed along the vessel wall may lead to progressive degeneration of the aortic media, increasing the likelihood of elastic fiber rupture and gradually reducing its thickness, while increasing the interlaminar spacing, thus promoting damage to the vessel wall tissue and inducing dilation. Under six rounded corner radius conditions, most of the wall surface was at a low to medium shear level; areas with higher WSS were mostly located in localized acceleration areas such as the outer side of the tortuous segment and near the branch openings.

[0216] Please see Figure 28 and Figure 29 , Figure 28 The image shown is a cloud map illustrating the wall shear stress distribution of heart valves with different fillet radii at peak systole within the same aorta in one embodiment of this application. Figure 29The figure shows a schematic diagram of the maximum wall shear stress at the peak contraction time for valves with different fillet radii in one embodiment of this application. As shown, the effect of fillet radius on the maximum wall shear stress is not monotonic: the maximum WSS for the first candidate leaflet model A is about 44 Pa, while that for the first candidate leaflet model B rises to 50.4 Pa; when the radius increases to the first candidate leaflet models C and D, the maximum WSS drops to 40.3 Pa and 40.7 Pa, respectively, which are the lowest in the group; when the radius continues to increase to the first candidate leaflet model E, the maximum WSS rises to the highest of 53.4 Pa, and then falls back to 45.8 Pa in the first candidate leaflet model F. Therefore, this figure shows that the fillet radius can change the spatial distribution and extreme value of the wall shear stress during the peak contraction period, with a medium fillet radius tending to reduce the maximum WSS and improve the uniformity of wall load.

[0217] To quantify the variation of wall shear stress with different corner radii, the average wall shear stress at typical locations on the aortic wall during peak systole was measured. (See also...) Figure 30 The figure shows a schematic diagram of the average wall shear stress of each model in a segmented region of the aorta in one embodiment of this application. As shown in the figure, the wall shear stress variation trend of the first candidate leaflet model A to the first candidate leaflet model F is generally consistent, showing a gradual increase from the proximal end, reaching a peak in the middle segment, and then a significant decrease in the distal end. There are certain differences between different types of valves. The first candidate leaflet models B and C are generally at a higher level, followed by the first candidate leaflet models D and E, with the first candidate leaflet model A in the middle. The first candidate leaflet model F shows a lower wall shear stress in most segmented regions. The results indicate that the first candidate leaflet model F has a lower wall shear load compared to other types, and its near-wall flow environment is more gentle.

[0218] In step S130, the radius of the free edge radius of the target leaflet is determined based on the multi-index simulation evaluation results of multiple first candidate leaflet models.

[0219] In one embodiment, determining the radius of the free edge of the target leaflet includes: selecting first candidate leaflet models whose opening performance evaluation results meet the first preset opening conditions based on the multi-index simulation evaluation results of multiple first candidate leaflet models, and determining the radius of the free edge of the target leaflet among the selected first candidate leaflet models based on the structural mechanics evaluation results and the hemodynamics evaluation results.

[0220] The preset opening conditions refer to the conditions used to select candidate leaflet models based on the opening performance evaluation results. Preset opening conditions may include: the geometric opening area of ​​the candidate leaflet model reaches a preset opening area threshold; the geometric opening area of ​​the candidate leaflet model is not less than a preset proportion among multiple candidate leaflet models; the geometric opening area of ​​the candidate leaflet model is within a preset ranking range among multiple candidate leaflet models; or the geometric opening area of ​​the candidate leaflet model meets other preset requirements related to valve opening performance. In some embodiments, the preset opening conditions may include a first preset opening condition and a second preset opening condition, described later. The first preset opening condition is used to select a first candidate leaflet model, and the second preset opening condition is used to select a second candidate leaflet model.

[0221] In one specific implementation, a preset opening area threshold can be set based on experience. The geometric opening area of ​​all candidate leaflet models obtained based on the method described above is compared with the preset opening area threshold, and one or more candidate leaflet models whose geometric opening area is greater than the preset opening area threshold are selected as the first candidate leaflet models.

[0222] Based on this, in one example, among the first candidate leaflet models selected, the leaflet free edge fillet radius corresponding to the first candidate leaflet model that satisfies the first preset structural conditions and the first preset flow field conditions is determined as the target leaflet free edge fillet radius.

[0223] The preset structural conditions refer to the conditions used to select candidate leaflet models based on the structural mechanics evaluation results. Preset structural conditions may include: at least one of the maximum principal stress, maximum principal strain, von Mises equivalent stress, and maximum shear stress of the candidate leaflet model is below a preset threshold; the stress and / or strain level of the candidate leaflet model is lower than the stress and / or strain level of at least one other candidate leaflet model; the structural mechanics indicators of the candidate leaflet model meet preset ranking requirements; or the structural mechanics indicators of the candidate leaflet model meet preset structural safety conditions. Preset structural conditions may include a first preset structural condition and a second preset structural condition, described later. The first preset structural condition is used to select the target leaflet free edge fillet radius from the first candidate leaflet models, and the second preset structural condition is used to select the target leaflet thickness from the second candidate leaflet models.

[0224] In one example, the first preset structural condition includes: the stress and / or strain level characterized by the structural mechanics evaluation results of the first candidate leaflet model determined to be the target leaflet free edge fillet radius is lower than the stress and / or strain level of at least one first candidate leaflet model with a smaller leaflet free edge fillet radius. This condition is used to characterize the effect of the target leaflet free edge fillet radius in reducing local stress concentration or improving structural response relative to a smaller fillet radius.

[0225] The preset flow field conditions refer to the conditions used to screen candidate leaflet models based on hemodynamic evaluation results. Preset flow field conditions may include: at least one of blood flow velocity, wall pressure, and wall shear stress meeting a preset threshold; the distribution of blood flow velocity, wall pressure, and / or wall shear stress meeting a preset uniformity requirement; the range of local high-velocity jets meeting a preset requirement; wall shear stress lower than the wall shear stress corresponding to one or more other candidate leaflet models; or hemodynamic indices meeting a preset ranking requirement. Preset flow field conditions may include a first preset flow field condition and a second preset flow field condition, described later. The first preset flow field condition is used to screen the target leaflet free edge radius from the first candidate leaflet models, and the second preset flow field condition is used to screen the target leaflet thickness from the second candidate leaflet models.

[0226] The first preset flow field conditions may include: the blood flow velocity distribution, wall pressure distribution, and / or wall shear stress distribution corresponding to the first candidate leaflet model meeting preset ranges, preset uniformity requirements, or preset relative comparison conditions. The first preset flow field conditions are used to evaluate the influence of different free edge fillet radii on the aortic flow field state when determining the target leaflet free edge fillet radius. In this embodiment, the geometric opening areas of all six first candidate leaflet models meet the first preset opening conditions, and therefore all enter the structural mechanics evaluation and hemodynamics evaluation stages. During parameter determination, the maximum wall shear stress is not used as the sole criterion; instead, the geometric opening area, the peak and spatial average values ​​of each structural mechanics index, the blood flow velocity distribution, the wall pressure distribution, and the average wall shear stress in each segment of the aorta are comprehensively considered. The first candidate leaflet model F has the largest geometric opening area, the lowest number of structural mechanics indices, and exhibits low average wall shear stress and relatively smooth velocity and pressure distributions in most aortic segment regions; therefore, its corresponding 3.5 mm is determined as the target leaflet free edge fillet radius. Although the maximum wall shear stress values ​​of some medium corner radius models are lower than those of the first candidate leaflet model F, this local extreme value is not used as a basis for rejecting the comprehensive evaluation result of the first candidate leaflet model F.

[0227] In one specific embodiment, this application compares the opening and closing behavior, mechanical response, and peak systolic hemodynamic characteristics of six aortic valve finite element models with different fillet radii during a cardiac cycle through finite element structural simulation and CFD numerical simulation based on the geometry of the aortic arch of a real patient, revealing the influence of fillet radius on valve function and durability. For the first candidate leaflet models A to F with different fillet radii of the leaflet free edge, each leaflet model can effectively open and close, with consistent deformation and displacement distribution patterns, and the maximum deformation is concentrated in the middle region of the free edge. Increasing the fillet radius makes the leaflet transition smoother and the opening and closing shape more regular. The GOA (Gross Opening Area) increases monotonically with increasing fillet radius, with the first candidate leaflet model F having the largest opening area, 14.9% higher than the first candidate leaflet model A, and the opening effect of the first candidate leaflet model F is the best. Increasing the fillet radius can significantly improve stress concentration, and there is a significant negative correlation between key mechanics and fillet radius. With increasing fillet radius, the decreases in maximum principal stress, maximum equivalent stress, maximum shear stress, and maximum principal strain are 28.0%, 18.0%, 24.7%, and 18.8%, respectively, indicating that increasing the fillet radius can effectively reduce the risk of stress concentration, fracture, and fatigue failure, and improve valve structural stability. Increasing the fillet radius can reduce flow velocity and shear force levels, and the effect of fillet radius on maximum blood flow velocity and maximum wall shear stress is non-monotonic. Although the first candidate leaflet model F is not the lowest in all local extreme values, it exhibits lower mean wall shear stress in most aortic segment regions and has a smoother velocity and wall pressure distribution. Considering its larger geometric opening area and lower structural stress and strain levels, the first candidate leaflet model F exhibits superior overall performance. Therefore, the determined target leaflet free edge fillet radius is 3.5 mm, and the first candidate leaflet model is F.

[0228] In step S140, based on the leaflet geometric model with the radius of the free edge radius of the target leaflet, multiple different leaflet thicknesses are assigned to generate multiple second candidate leaflet models. The obtained second candidate leaflet models are represented in the form of second candidate leaflet geometric models.

[0229] In one embodiment, multiple second candidate leaflet models all have the target leaflet free edge fillet radius and the same leaflet geometric topology. In a specific example, the leaflet geometric model with the target leaflet free edge fillet radius is the first candidate leaflet model F with a leaflet free edge fillet radius of 3.5 mm.

[0230] In one embodiment, the multiple different leaflet thicknesses are selected from a range of values ​​within the range of 0.4 mm to 0.6 mm. The leaflet thickness is one of the important design parameters affecting the overall performance of the artificial aortic valve. While maintaining a basically consistent leaflet geometry, multiple second candidate leaflet models are constructed by changing the leaflet thickness to analyze the impact of different thicknesses on valve opening and closing performance, structural mechanical properties, and hemodynamic performance, and to determine the target leaflet thickness that meets the comprehensive evaluation requirements.

[0231] In one example, the 0.4 mm to 0.6 mm range covers relatively thin, medium, and relatively thick leaflet designs. Thinner leaflets typically have lower flexural stiffness, opening more easily under transvalvular pressure drop, resulting in a larger effective opening area and lower transvalvular pressure drop, but may increase the risk of localized stress, strain, and fatigue damage. Thicker leaflets have higher structural stiffness, helping to reduce leaflet deformation and localized stress concentration, improving structural stability and durability, but may reduce leaflet opening amplitude and increase flow resistance and energy loss. Therefore, designing multiple candidate thicknesses within the aforementioned range is beneficial for achieving a comprehensive balance between valve opening performance, structural safety, and hemodynamic performance.

[0232] In one example, the multiple different leaflet thicknesses include at least three of 0.4 mm, 0.45 mm, 0.5 mm, 0.55 mm, and 0.6 mm. For example, multiple second candidate leaflet models corresponding to different thicknesses can be established, and while keeping the leaflet profile curve, leaflet height, free edge shape, connection boundary, and other structural parameters consistent, only the leaflet thickness is changed as a single design variable to reduce the influence of other factors on the evaluation results and improve the comparability between different thickness schemes.

[0233] Please see Figure 31 The figure shows a top view of five sets of tricuspid artificial aortic valves with different leaflet thicknesses in one embodiment of this application. As shown in the figure, the thicknesses of the five models are 0.4 mm, 0.45 mm, 0.5 mm, 0.55 mm and 0.6 mm, respectively. Based on the first candidate leaflet model F with a leaflet free edge radius of 3.5 mm, five second candidate leaflet models are generated for the above five different leaflet thicknesses. They are respectively called second candidate leaflet model A (marked as Type A in the figure), second candidate leaflet model B (marked as Type B in the figure), second candidate leaflet model C (marked as Type C in the figure), second candidate leaflet model D (marked as Type D in the figure) and second candidate leaflet model E (marked as Type E in the figure).

[0234] In one embodiment, the leaflet material in the second candidate leaflet model is characterized using a hyperelastic constitutive model. In one example, the hyperelastic constitutive model is a second-order Ogden constitutive model. The method for obtaining the second candidate leaflet model is the same as or similar to the method for obtaining the first candidate leaflet model described above, and will not be repeated here.

[0235] In step S150, the multi-index simulation evaluation is performed on multiple second candidate leaflet models.

[0236] The second candidate leaflet model for multi-index simulation evaluation is represented as a second candidate leaflet finite element model. Therefore, before performing multi-index simulation evaluation on multiple second candidate leaflet models, a step is included to construct a second candidate leaflet finite element model based on each second candidate leaflet geometric model. Specifically, each second candidate leaflet geometric model is assigned the same material properties, a finite element mesh is generated, and the contact relationship between the leaflets, analysis step, inter-leaf pressure differential load, and boundary constraints are set to obtain the second candidate leaflet finite element model. Thus, each second candidate leaflet finite element model uses the same target leaflet free edge fillet radius, material parameters, finite element modeling rules, load conditions, and boundary conditions, but has different leaflet thicknesses, so that the differences in the finite element analysis results are mainly attributed to variations in leaflet thickness.

[0237] In one embodiment, the multi-index simulation evaluation includes the opening performance evaluation, the structural mechanics evaluation, and the hemodynamic evaluation.

[0238] Please see Figure 32 The figure shows a schematic diagram of the dynamic changes of the second candidate leaflet model at different times in one embodiment of this application. As shown, it illustrates the deformation characteristics of the valves from second candidate leaflet model A to second candidate leaflet model E during the dynamic process. Each group of models underwent a complete opening and closing process, but the difference in valve thickness had a significant impact on the opening amplitude and deformation coordination. The thinner models (second candidate leaflet model A and second candidate leaflet model B) exhibited more complete valve opening and a larger opening profile during the opening phase, with a more obvious outward expansion trend at intermediate moments, indicating higher compliance and greater sensitivity to fluid load response. As the thickness increased to second candidate leaflet model C, second candidate leaflet model D, and second candidate leaflet model E, the valve stiffness gradually increased, the outward expansion during opening weakened, and the opening profile relatively contracted, indicating that although the thicker valve has better structural stability, its motion flexibility decreased. During the closing phase, each model gradually forms a typical trilobed closure shape. The thicker model has a more regular closed contour, reflecting its greater stability in deformation control and structural support, while the thinner model is more prone to larger transient deformations.

[0239] Insufficient deformation during the initial valve opening phase can adversely affect cardiac function. Comparing the dynamic deformation characteristics of five models with different thicknesses, the second candidate leaflet model C performed best overall. Compared to second candidate leaflet models A and B, second candidate leaflet model C did not exhibit excessive local deformation, demonstrating better structural stability. Compared to second candidate leaflet models D and E, second candidate leaflet model C also did not show significant opening restriction, maintaining good valve mobility and a relatively regular opening and closing process. Therefore, second candidate leaflet model C achieves a better balance between flexibility, opening capacity, and closing stability, and can be considered the type with the best overall performance among the five thickness schemes.

[0240] In one embodiment, the apical performance evaluation includes determining the geometric opening area of ​​the second candidate leaflet model in the open state; the geometric opening area is the valve orifice opening area formed when the second candidate leaflet model reaches its maximum opening degree during the cardiac cycle.

[0241] The geometric opening area is calculated directly from the deformed valve geometry. Quantification of the geometric opening area of ​​different models shows that the valve geometric opening area decreases with increasing valve thickness, but the geometric opening areas of all five second candidate leaflet models meet the preset second opening condition of this embodiment.

[0242] In one embodiment, the structural mechanical evaluation includes: applying a transvalvular pressure differential load during the cardiac cycle to the second candidate leaflet model, and performing finite element dynamic opening and closing analysis on the second candidate leaflet model.

[0243] In one embodiment, the structural mechanics evaluation includes determining the maximum principal stress, maximum principal strain, von Mises equivalent stress, and maximum shear stress of the second candidate leaflet model during dynamic opening and closing.

[0244] Please see Figure 33a and 33b , Figure 33a The graph shown is a curve illustrating the change of maximum principal stress on the valve over time in one embodiment of this application. Figure 33bThe figure shows the curve of the average maximum principal stress on the valve versus time in one embodiment of this application. As shown, the maximum principal stress of valves of different thicknesses fluctuates dynamically over time, with varying amplitudes and peak values. Overall, as the valve thickness increases from 0.4 mm (second candidate leaflet model A) to 0.6 mm (second candidate leaflet model E), the maximum principal stress and the average maximum principal stress generally show a decreasing trend, indicating that increasing the thickness helps improve the valve's stress state and reduce local stress concentration. In the early stage of the cycle, especially before 0.2 s, the thinner valves exhibit more significant stress fluctuations, which is related to their lower bending stiffness and greater susceptibility to deformation under transvalvular pressure differential and transient flow field. In contrast, the thicker valves have a more stable stress response due to their higher structural stiffness. As the motion progresses, the stress fluctuations of all models decrease, but the thinner valves still maintain a higher stress level overall, indicating that they bear a greater mechanical load throughout the cardiac cycle. The variation pattern of the average maximum principal stress is consistent with that of the instantaneous maximum principal stress, i.e., the second candidate leaflet model A has the highest stress, while the second candidate leaflet model E has the lowest stress. This further indicates that increasing the thickness not only reduces local peak stress but also helps improve the overall stress uniformity. Therefore, it is suggested that although thinner valves have higher flexibility, they are more prone to larger stress fluctuations and peak values, which increases the risk of fatigue damage. Thicker valves have advantages in reducing stress and improving dynamic stability, but excessive thickness may adversely affect valve opening performance.

[0245] Please see Figure 34 The figure shows the maximum principal stress cloud map of the valve at different times for five groups of second candidate leaflet models in one embodiment of this application. As shown in the figure, the distribution pattern of high stress concentration areas in each model is basically consistent, mainly located in the ventral region of the valve and the central region of the free edge. The maximum principal stress distribution characteristics are closely related to the valve's stress mechanism: the ventral region is the main load area bearing blood flow pressure during diastole, while the central region of the free edge bears both pressure and traction effects during opening and closing, thus making it more prone to stress concentration. This result is also consistent with clinical observations that fatigue damage of artificial biological valves often occurs in the above-mentioned areas.

[0246] Table 8 lists the peak and average values ​​of the maximum principal stresses for all candidate second-stage valve models. There are some differences in the maximum principal stresses among the five candidate models, with candidate model A showing a significantly higher peak value. Quantitative calculations show that the maximum principal stress of candidate model A is 8.87%, 22.68%, 28.75%, and 29.17% higher than that of candidate models B, C, D, and E, respectively. This significantly increased stress level indicates that the valve in candidate model A needs to continuously withstand higher mechanical loads during the menstrual cycle, significantly increasing the risk of material fatigue damage. The peak values ​​of the maximum principal stresses for candidate models C, D, and E are relatively similar. The stress values ​​of candidate models D and E are almost equal, while candidate model C is only 0.109 MPa higher than candidate model E. This indicates that the stress-bearing state of these three models is relatively stable and within a lower stress level range.

[0247] As valve thickness increases, the stress concentration effect gradually weakens, and the strength of high-stress areas decreases significantly. This indicates that thicker valves have a stronger load-dispersing ability and can reduce local stress accumulation. Since long-term stress concentration accelerates material fatigue aging, the lower and more stable stress levels of the second candidate leaflet models C, D, and E are more conducive to reducing the risk of fatigue damage and improving valve durability.

[0248] Table 8 Maximum principal stress of the aortic valve

[0249] Please see Figure 35a and 35b , Figure 35a The graph shown is a curve illustrating the maximum principal strain on the valve over time in one embodiment of this application. Figure 35b The figure shows a curve illustrating the variation of the average maximum principal strain with time in one embodiment of this application. As shown, both the maximum principal strain and the average maximum principal strain of the valve exhibit significant periodic changes over time, with significant differences between models of different thicknesses. As the valve thickness increases from the second candidate leaflet model A (0.4 mm) to the second candidate leaflet model E (0.6 mm), the maximum principal strain generally shows a decreasing trend, indicating that thickening can effectively suppress tensile deformation during dynamic opening and closing. Figure 35aThe results show that the thinner models (second candidate leaflet model A and second candidate leaflet model B) have higher strain peaks and more obvious fluctuations in the early stage of the cycle, indicating that they are more prone to local overstretching under transient pressure difference. As the thickness increases, the strain levels of the second candidate leaflet model C, the second candidate leaflet model D, and the second candidate leaflet model E gradually decrease, and the fluctuations also tend to be smoother, indicating that the thicker valve can effectively limit local elongation and reduce strain concentration due to its higher structural stiffness. Figure 35b The variation law of the average maximum principal strain is basically consistent with the instantaneous results. The average strain of each model gradually decreases with the increase of thickness, and the curve of the model with larger thickness is more stable, indicating that the thickening not only reduces the local extreme tensile deformation, but also improves the overall deformation coordination of the valve.

[0250] Please see Figure 36 The figure shows the maximum principal strain cloud maps of five groups of second candidate leaflet models at different times in one embodiment of this application. As shown, the maximum principal strain distribution characteristics of second candidate leaflet models A to E at four key moments in the cardiac cycle are illustrated. During the valve opening phase, the strain distribution generally exhibits a lower and more uniform distribution at the bottom, with concentrations at the edges and abdomen. The bottom of the valve, being fixedly connected to the valve annulus and having a smaller range of motion, consistently exhibits a lower and more uniform strain level. The free edge, as a non-fixed end, requires a larger expansion during opening, while the abdominal region bears the main blood flow load; therefore, both are more prone to strain concentration. As the valve enters the closing phase, the strain concentration area shifts. The concentration at the free edge gradually weakens, while the fusion area and abdominal region become the main high-strain areas. This change indicates that during the closing process, influenced by reverse pressure loads and the mutual squeezing and pulling between the leaflets, the deformation accumulation at the fusion area and the abdomen is more significant, thus becoming new strain concentration areas.

[0251] As shown in Table 9, the peak value and average maximum principal strain generally decrease with increasing valve thickness. The peak principal strain decreases significantly from the second candidate leaflet model A to the second candidate leaflet model C, and tends to stabilize after the second candidate leaflet model C, while the average principal strain continues to decrease. Thickening can effectively suppress local excessive elongation during the closure process, but when the thickness increases to 0.5 mm, its improvement on local peak deformation shows a marginal weakening trend.

[0252] Table 9 Maximum principal strain of the aortic valve

[0253] Please refer to Figure 37a and Figure 37b , Figure 37a The graph shown is a curve illustrating the change of maximum equivalent stress on the valve over time in one embodiment of this application. Figure 37bThe figure shows a curve illustrating the variation of average maximum equivalent stress over time in one embodiment of this application. As shown, the maximum equivalent stress and average maximum equivalent stress of different thickness models exhibit significant dynamic changes throughout the entire period. Overall, as the valve thickness increases from the second candidate leaflet model A to the second candidate leaflet model E, both show a decreasing trend, indicating that thickening helps alleviate stress concentration in the valve during dynamic movement. Figure 37a The results show that the maximum equivalent stress of each model fluctuates periodically over time, with significant peaks appearing in the early and middle stages of the cycle. The second candidate leaflet model A exhibits the highest stress level and the most significant fluctuations, indicating that thinner valves, due to their lower stiffness, are more prone to large local deformations under transient loads, resulting in higher equivalent stress and stronger stress oscillations. As the thickness increases, the stress peak values ​​of the second candidate leaflet models B to E gradually decrease, and the curves become flatter, indicating that thicker valves have stronger load dispersion capabilities and can more effectively suppress local stress concentration. Figure 37b The variation law of the average maximum equivalent stress is consistent with the instantaneous results. The average stress of each model gradually decreases with increasing thickness, with the second candidate leaflet model A having the highest stress and the second candidate leaflet model E having the lowest stress. Moreover, the thicker models show smaller fluctuations throughout the cycle. The results indicate that thickening not only reduces the overall stress level of the valve but also helps to improve its mechanical stability and durability under cyclic loading.

[0254] Please see Figure 38 The figure shows the equivalent stress cloud map of the valve at different times for five groups of second candidate leaflet models in one embodiment of this application. As shown, it illustrates the equivalent stress distribution characteristics of the five groups of second candidate leaflet models (A to E) during a complete cardiac cycle. Overall, the equivalent stress distribution patterns of each model are basically consistent, with the main differences being the phased migration of stress concentration areas and the magnitude of peak values. During the valve opening phase, high equivalent stress is mainly concentrated in the abdominal and free edge regions. This is related to the fact that the abdominal region bears the main positive blood flow load and the free edge undergoes significant movement during this phase. As the cardiac cycle enters the closing phase, the stress concentration area gradually migrates from the abdominal region to the fusion point. This is because during the closing process, the valve needs to withstand reverse blood flow impact, and the fusion point, as a key connection between the leaflets, gradually becomes the main load-bearing area, thus exhibiting more obvious stress concentration.

[0255] Table 10 shows the quantitative results of peak equivalent stress, which further reveals the relationship between valve thickness and stress load: the maximum equivalent stress of the five candidate leaflet models shows a continuous decreasing trend with increasing thickness. Quantitative comparison reveals that the maximum equivalent stress of candidate leaflet model A is significantly higher than all other models, increasing by 7.46%, 29.74%, 23.2%, and 30.57% compared to candidate leaflet models B, C, D, and E, respectively. This indicates that an excessively thin valve structure must bear an overall stress load far exceeding normal levels. The maximum equivalent stress of candidate leaflet models C, D, and E shows small differences, with the difference between candidate leaflet models D and E being only 0.089 MPa, indicating that the equivalent stress level tends to stabilize after the valve thickness increases to a certain range.

[0256] Table 10 Maximum equivalent stress of the aortic valve

[0257] From a mechanical perspective, increasing valve thickness enhances its overall load-bearing capacity, distributing the mechanical load during the cardiac cycle over a wider area, thereby reducing local stress concentration and lowering the overall stress level. This reduction in stress load is crucial for improving valve durability. This result corroborates the aforementioned maximum principal stress analysis: the maximum principal stress reflects the risk of localized concentrated loads, while the equivalent stress explains the role of thickness optimization in reducing mechanical load and improving structural stability at the overall level, together providing a mechanical basis for optimizing the thickness parameters of artificial bioprosthetic valves.

[0258] Please see Figure 39a and 39b , Figure 39a The graph shown is a curve illustrating the change of maximum shear stress on the valve over time in one embodiment of this application. Figure 39b The figure shows a curve illustrating the variation of the average maximum shear stress on the valve over time in one embodiment of this application. As shown, the maximum shear stress and average maximum shear stress of each thickness model also fluctuate periodically over time, but generally decrease gradually with increasing thickness. Thinner valves exhibit more pronounced shear stress fluctuations in the early and peak stages of the cycle, indicating that they are more prone to strong tangential stress responses during opening and closing due to bending, torsion, and local contact incoordination. Thicker valves, on the other hand, have increased structural stiffness and enhanced shear deformation resistance, resulting in smoother load transfer and mitigation of local shear concentration. Therefore, from the perspective of maximum shear stress, thickening can effectively improve the tangential stress state of the valve during dynamic opening and closing, enhance the structural coordination deformation capacity, and reduce the potential damage risk caused by repeated shearing.

[0259] Please see Figure 40 The figure shows the maximum shear stress cloud map of the valve at different times for five groups of second candidate leaflet models in one embodiment of this application. The figure illustrates the maximum shear stress distribution characteristics of second candidate leaflet models A to E. The maximum shear stress distribution patterns of the second candidate leaflet models with different thicknesses are highly consistent with the aforementioned equivalent stress distribution, indicating that under cardiac cycle loads, shear stress and equivalent stress have similar transmission paths, and their concentrated areas correspond to the main load-bearing parts of the valve, thus further supporting the stability and reliability of the mechanical simulation results of this application. During the valve opening phase, the abdominal region, while bearing positive blood flow pressure, is also affected by blood flow shearing and valve bending deformation, resulting in significant shear stress concentration. During the valve closing phase, the fusion site also exhibits a high level of shear stress under reverse impact loads. This result indicates that the abdominal region and the fusion site are not only the main areas affected by tensile and compressive loads but also sensitive areas to shear loads, and should be given special attention in valve structure design.

[0260] As shown in Table 11, the average maximum shear stress of the valve generally decreases with increasing thickness, indicating that thickening helps improve the overall shear stress state during closure. The second candidate leaflet models C and E have lower local shear peak values, while the second candidate leaflet model D shows a slight increase, suggesting that the local peak value, in addition to being affected by thickness, may also be related to the contact state and deformation coordination at the moment of closure. In contrast, the maximum shear stress of the second candidate leaflet model A is significantly higher than the other models, indicating that an excessively thin valve is more susceptible to higher tangential loads during dynamic opening and closing, thus increasing the risk of shear damage to the material. In summary, increasing the valve thickness can effectively reduce tangential stress concentration during the closure stage and improve structural stability. The second candidate leaflet model E performs better in terms of overall shear stress control, while the second candidate leaflet model C is more advantageous in suppressing local peak values.

[0261] Table 11 Maximum shear stress of the aortic valve

[0262] To better analyze the impact of different valve thicknesses on the magnitude of valve mechanical parameters, in one specific implementation, multiple sample points were randomly collected from the valve region regarding maximum principal stress, von Mises equivalent stress, maximum shear stress, and maximum principal strain. 54 sample points were selected for each model, resulting in a sample size of 270. Based on valve thickness levels of 1, 2, 3, 4, and 5, Spearman correlation analysis was performed in Graphpad Prism 9.1.0. The results showed a significant monotonically negative correlation between valve thickness and all mechanical response indices (P < 0.0001). The strongest negative correlation was observed between valve thickness and principal stress (r = -0.814), indicating that principal stress is most sensitive to changes in valve thickness. Strong negative correlations were also observed between valve thickness and principal strain (r = -0.662), shear stress (r = -0.668), and equivalent stress (r = -0.649). Overall, as the valve thickness increases, all mechanical properties show a decreasing trend, indicating that appropriately increasing the valve thickness may help improve the valve's stress state and enhance its mechanical stability. Combining the results of valve dynamic changes, geometric opening area, and mechanical response, it is evident that there is a clear trade-off between opening performance and structural safety for valves of different thicknesses. While thinner models have larger opening areas, they also exhibit higher stress-strain levels and more significant dynamic fluctuations; thicker models, while effectively reducing dangerous stress and deformation, may limit valve opening to some extent. In contrast, the second candidate leaflet model C demonstrates a better balance between geometric opening area, stress-strain control, and dynamic stability, and can therefore be considered the valve thickness with the best overall performance in this application.

[0263] In one embodiment, the hemodynamic evaluation includes determining the blood flow velocity, wall pressure, and wall shear stress corresponding to the second candidate leaflet model after it is placed into the aortic flow field model.

[0264] Please see Figure 41 The figure shows the blood velocity streamline distribution cloud map of each thickness model during the peak systolic phase in one embodiment of this application. As shown, at the peak systolic phase, each thickness model exhibits a similar overall flow pattern: the high-velocity zone is mainly located at the valve outlet and the proximal ascending aorta, and gradually diffuses and attenuates along the path; the aortic arch exhibits eccentric wall-attached flow due to the tortuous effect. The main differences between different thicknesses lie in the jet intensity, concentration, and continuity along the aortic arch.

[0265] As the valve thickness increased from 0.40 mm to 0.50 mm, the high-velocity region near the valve orifice weakened overall. The high-velocity core diffused more easily in the proximal ascending aorta, and the continuity of the high-velocity band decreased at the aortic arch and the aortic inlet, manifested as a reduction in the size of the high-velocity region and an increase in the proportion of the low-velocity region. Specifically, the jet distribution in the second candidate leaflet model C was more dispersed, indicating a weakening of the local velocity extreme at the peak. In contrast, when the thickness was further increased to 0.55 mm (second candidate leaflet model D), a more concentrated high-velocity core formed at the valve orifice, the high-velocity region expanded, and the high-velocity band extended more significantly towards the aortic arch, indicating enhanced jet convergence and a larger local velocity gradient. When the thickness was further increased to 0.60 mm (second candidate leaflet model E), the high-velocity region remained significant, but weakened compared to the second candidate leaflet model D, with its jet intensity and concentration falling between those of the second candidate leaflet model D and the thinner models.

[0266] Please see Figure 42 The figure shows a schematic diagram illustrating the maximum blood velocity of different thickness models during the peak contraction phase in one embodiment of this application. As shown, the maximum velocity of different thickness models exhibits a non-monotonic variation, being lowest at 0.50 mm and highest at 0.55 mm. This result indicates that thickness mainly affects the jet convergence by influencing the peak valve opening morphology and effective opening area, thereby altering the local velocity extreme value. Specifically, the 0.55 mm model has the most concentrated jet and the highest velocity, the 0.50 mm model has the most dispersed jet and the lowest velocity, while the 0.60 mm model shows a slight decrease due to factors such as limited opening.

[0267] To quantify blood flow velocity variations in models of different thicknesses, the maximum and average flow velocities at typical locations in the aorta during peak systole were measured. Please refer to [link / reference]. Figure 43a and 43b , Figure 43a This diagram illustrates the maximum blood flow velocity in a segmented aortic region according to one embodiment of this application. Figure 43b The figure shows a schematic diagram of the average blood flow velocity in the aortic region in one embodiment of this application. As shown, the overall trends of the second candidate leaflet models A to E are similar, but there are still some differences in the maximum blood flow velocity. Specifically, the blood flow velocity trends along the aortic region are similar for second candidate leaflet models A to E, but there are still differences in the local maximum velocity. Combined with... Figure 42 and Figure 43a It can be seen that the velocity extrema and maximum velocity in some segmented regions of the second candidate leaflet model D are relatively high, followed by the second candidate leaflet models E and A; the velocity extrema of the second candidate leaflet models B and C are relatively low, with the second candidate leaflet model C having the lowest global maximum velocity. In comparison, the average blood flow velocity differences among the models are relatively small, indicating that thickness variations mainly affect the local velocity extrema and the degree of jet concentration.

[0268] Please see Figure 44 The figure shows a cross-sectional cloud map of the intraluminal velocity of the aorta during the peak systolic phase in one embodiment of this application. As shown, it illustrates the cross-sectional distribution of blood flow velocity in the same aorta for second candidate leaflet models of different thicknesses during the peak systolic phase. The results show that second candidate leaflet models A to E all exhibit varying degrees of eccentric flow, with the high-velocity region mainly concentrated on one side of the cross-section and extending downstream along the blood flow direction. Some differences remain between the different models. Second candidate leaflet model D exhibits a more pronounced eccentric high-velocity jet characteristic, with a clear and continuous high-velocity core region visible in sections A to C, indicating that it is more likely to form a strong and stable local high-velocity blood flow. Second candidate leaflet models A and B are similar, also exhibiting significant eccentric flow, but the continuity and concentration of the high-velocity region are slightly weaker than in second candidate leaflet model D; second candidate leaflet model E is next; and second candidate leaflet model C is generally in the middle.

[0269] Please see Figure 45a and 45b , Figure 45a The image shown is a pressure distribution cloud map of the aortic wall under the same conditions for models of different thicknesses in one embodiment of this application. Figure 45b This is shown as the maximum wall pressure under the same conditions for models of different thicknesses in one embodiment of this application. Figure 45a It is evident that the overall pressure level of each model is predominantly moderate, but localized high-pressure zones appear in the bending transition section and areas with significant local geometric changes. Models of different thicknesses exhibit certain differences in pressure distribution characteristics. Specifically, the second candidate leaflet model D and the second candidate leaflet model E show more pronounced high-pressure zones in the lower section of the structure, indicating significant local pressure concentration in these areas. The second candidate leaflet model A also shows some high-pressure distribution in the middle section, but the range is relatively small. In contrast, the second candidate leaflet model B and the second candidate leaflet model C show more uniform pressure distribution and smaller high-pressure zones, indicating a relatively weaker local pressure concentration and a more stable stress state on the wall surface.

[0270] Combination Figure 45bThe maximum pressure results show that the differences between models with different thicknesses can be further quantified. The order of magnitude of the maximum pressure for the second candidate leaflet models A, B, C, D, and E is: D > E > A > B > C. Among them, the second candidate leaflet model D has the highest maximum pressure, indicating that the local wall surface experiences the most significant pressure load under this thickness condition. The second candidate leaflet models E and A are next, still exhibiting significant local pressure concentration. In contrast, the maximum pressure of the second candidate leaflet model B is lower, while that of the second candidate leaflet model C is the lowest, indicating that its peak pressure level is the smallest and its local pressure state is relatively more moderate.

[0271] To quantify the wall pressure changes in different models, the average wall pressure at typical locations on the aortic wall during peak systole was measured. Please refer to [link / reference needed]. Figure 46 The figure shows a schematic diagram of the average wall pressure in a segmented region of the aorta in one embodiment of this application. As shown, the wall pressure variation trends of the second candidate leaflet models A to E in typical aortic regions are generally consistent, indicating that different models have not changed the basic law of wall pressure distribution along the aorta, but there are still slight differences in local peak values. Regarding the average wall pressure, the curves of each type are more similar, generally increasing gradually from S1 to S5, and basically converging at S4 and S5. Overall, the difference in wall pressure between different types is significantly smaller than the difference in velocity field.

[0272] Please see Figure 47a and 47b , Figure 47a The image shown is a cloud map illustrating the shear stress distribution of the aortic wall under the same conditions for models of different thicknesses in one embodiment of this application. Figure 47b The figure shows the maximum wall shear stress (WSS) under the same conditions for models of different thicknesses in one embodiment of this application. The WSS reflects the tangential friction of the fluid on the wall. A localized increase in WSS usually indicates enhanced wall shear load, which may increase the risk of material fatigue or structural damage. Therefore, a lower and more uniformly distributed WSS is more beneficial to structural stability. Figure 47aIt is evident that the overall wall shear stress of each model is at a low level; higher WSS is mainly distributed in the bending section and local geometric transition areas. Models of different thicknesses show differences in local distribution. The second candidate leaflet model D and the second candidate leaflet model E exhibit more pronounced high WSS bands in the central region, indicating significant local shear stress concentration in this area; the second candidate leaflet model A also has a certain range of moderate stress zones. In contrast, the high WSS region of the second candidate leaflet model B is somewhat weakened, while the second candidate leaflet model C is predominantly low WSS, with only slight increases in a few local areas, indicating a more uniform wall shear stress distribution and a smaller range of high shear areas.

[0273] Further integration Figure 47b The maximum WSS values ​​in the data provide a more intuitive comparison of the differences between different thicknesses. The maximum WSS of the second candidate leaflet model A is approximately 61.3 Pa, that of the second candidate leaflet model B is approximately 49.8 Pa, that of the second candidate leaflet model C is approximately 44.5 Pa, that of the second candidate leaflet model D is approximately 65.8 Pa, and that of the second candidate leaflet model E is approximately 62.7 Pa. Based on these data, the maximum WSS values ​​follow the following order: second candidate leaflet model D > second candidate leaflet model E > second candidate leaflet model A > second candidate leaflet model B > second candidate leaflet model C. Among these, the second candidate leaflet model D has the highest maximum WSS, reaching 65.8 Pa, indicating that the scouring effect of the fluid on the wall is strongest under this thickness condition, and the local shear stress concentration phenomenon is also most obvious. The maximum WSS values ​​of the second candidate leaflet models E and A are also at a high level, indicating that their structures still have strong local shear effects. In contrast, the maximum WSS of the second candidate leaflet model B is significantly lower, while the maximum WSS of the second candidate leaflet model C is the lowest, at only 44.5 Pa, indicating that it produces the smallest peak wall shear stress among all the schemes.

[0274] Please see Figure 48The figure shows a schematic diagram of the average wall shear stress in a segmented region of the aorta in one embodiment of this application. As shown, the wall shear stress distribution pattern along the aorta is consistent for models of different thicknesses (second candidate leaflet models A to E) in a typical region: it gradually increases from the proximal to the middle segment, reaches a high level near S3 and S4, and then decreases significantly at S5. This indicates that the thickness change does not change the basic trend of shear stress distribution, but there are differences in local peak values ​​and amplitudes. Specifically, the second candidate leaflet model D has the highest shear stress at S2, indicating that its anterior and middle segment of the aorta bears a stronger average shear load; the second candidate leaflet model E is slightly higher at S3; the second candidate leaflet model C is generally lower in the S1 to S3 segments, and is close to other types at S4; the second candidate leaflet model B is the lowest at S5. In summary, the distribution patterns of the five models are similar, but the second candidate leaflet model D has a higher overall shear stress level, and the wall is more significantly subjected to shear; the second candidate leaflet model C is generally more gradual, especially the anterior segment with lower shear stress, and the local shear load is relatively weak.

[0275] Therefore, it can be concluded that the thickness scheme corresponding to the second candidate leaflet model C performs best among the five schemes. This scheme not only has the lowest maximum wall shear stress, but also the smallest high-stress region and a more uniform stress distribution, which can effectively reduce local shear load, thereby improving the stability and safety of the overall structure. Therefore, in this application, the thickness scheme corresponding to the second candidate leaflet model C is considered the optimal design scheme.

[0276] In step S160, the target leaflet thickness is determined based on the multi-index simulation evaluation results of multiple second candidate leaflet models, so as to obtain the target leaflet free edge radius and target leaflet thickness for the design of the trilobal artificial biological aortic valve.

[0277] In one embodiment, determining the target leaflet thickness includes: selecting second candidate leaflet models whose opening performance evaluation results meet the second preset opening conditions based on the multi-index simulation evaluation results of multiple second candidate leaflet models, and determining the target leaflet thickness among the selected second candidate leaflet models based on the structural mechanics evaluation results and hemodynamics evaluation results.

[0278] In one example, among the selected second candidate leaflet models, the leaflet thickness corresponding to the second candidate leaflet model that satisfies both the second preset structural condition and the second preset flow field condition is determined as the target leaflet thickness; wherein, the second preset structural condition includes: the stress and / or strain level characterized by the structural mechanical evaluation result of the second candidate leaflet model determined as the target leaflet thickness is lower than the stress and / or strain level of at least one second candidate leaflet model with a smaller leaflet thickness; the second preset flow field condition includes: the wall shear stress characterized by the hemodynamic evaluation result of the second candidate leaflet model determined as the target leaflet thickness satisfies the preset wall shear stress condition.

[0279] In this embodiment, the geometric opening areas of all five second candidate leaflet models satisfy the second preset opening condition. The leaflet thickness of second candidate leaflet model C is 0.5 mm, and its maximum principal stress, maximum principal strain, von Mises equivalent stress, and maximum shear stress are all lower than at least one second candidate leaflet model with a smaller leaflet thickness. Simultaneously, its maximum wall shear stress is approximately 44.5 Pa, lower than the approximately 49.8 Pa corresponding to second candidate leaflet model B with a smaller leaflet thickness, and lower than the approximately 65.8 Pa and approximately 62.7 Pa corresponding to second candidate leaflet models D and E with larger leaflet thicknesses, respectively. Therefore, second candidate leaflet model C simultaneously satisfies the second preset structural condition and the second preset flow field condition, and its corresponding 0.5 mm is determined as the target leaflet thickness.

[0280] The second preset flow field condition may further include: the wall shear stress corresponding to the second candidate leaflet model determined to be the target leaflet thickness is lower than that corresponding to at least one second candidate leaflet model with a leaflet thickness less than the target leaflet thickness and at least one second candidate leaflet model with a leaflet thickness greater than the target leaflet thickness.

[0281] The preset wall shear stress condition refers to the condition used to evaluate whether the wall shear stress corresponding to the second candidate leaflet model meets the hemodynamic design requirements. The preset wall shear stress condition may include: the wall shear stress is lower than a preset threshold; the wall shear stress is lower than the wall shear stress corresponding to at least one candidate model with a smaller thickness and at least one candidate model with a larger thickness; or the wall shear stress distribution meets a preset uniformity requirement. By setting the preset wall shear stress condition, it is possible to avoid selecting excessively thick leaflets solely based on the reduction in structural mechanical properties caused by increased thickness, thus ensuring that the target leaflet thickness balances structural stability and hemodynamic performance.

[0282] The stress and / or strain level refers to the index level used to reflect the degree of structural mechanical response of the candidate leaflet model. It can be characterized by the time history peak value, spatial average value, or a combination thereof of at least one of the following indices: maximum principal stress, maximum principal strain, von Mises equivalent stress, and maximum shear stress. A lower stress and / or strain level of the candidate leaflet model can indicate that at least one of its corresponding structural mechanical indices is lower than that of other candidate leaflet models, or that it meets the preset structural safety conditions.

[0283] In one example, the second preset flow field condition includes: the wall shear stress corresponding to the second candidate leaflet model determined to be the target leaflet thickness is lower than the wall shear stress corresponding to at least one second candidate leaflet model with a leaflet thickness less than the target leaflet thickness and at least one second candidate leaflet model with a leaflet thickness greater than the target leaflet thickness.

[0284] In one specific embodiment, this application analyzes the structural mechanical behavior and hemodynamic characteristics of artificial biological valves with different thicknesses (0.4–0.6 mm) during the cardiac cycle based on the finite element method and computational fluid dynamics methods, particularly the influence of valve thickness on valve opening and closing performance, geometric opening area, stress-strain distribution, and hemodynamic characteristics. For the second candidate leaflet models A to E with the target leaflet free edge radius, the overall deformation mode of valves with different thicknesses is basically consistent during dynamic opening and closing, with the maximum deformation mainly concentrated in the free edge region of the valve. As the valve thickness increases, the valve structural stiffness increases, and the leaflet deformation gradually decreases. A valve that is too thin (second candidate leaflet model A) exhibits significant deformation during the closure phase, easily leading to stress concentration. Conversely, a valve that is too thick (second candidate leaflet model D and subsequent models) has insufficient deformation during the initial opening phase due to increased stiffness and inertia, potentially affecting its normal opening performance. Second candidate leaflet models B and C have larger maximum opening areas, indicating stronger opening capabilities. However, as thickness increases further, the opening area gradually decreases, suggesting that excessive thickness limits the valve's opening range. With increasing valve thickness, the maximum principal stress, equivalent stress, maximum shear stress, and maximum principal strain all show an overall decreasing trend. Second candidate leaflet model A exhibits the highest stress and strain levels, while the stress levels of second candidate leaflet models C and subsequent models show relatively small differences, indicating that appropriately increasing thickness can effectively reduce stress concentration and improve structural stability. Different valve thicknesses have a certain impact on the blood flow structure within the aorta. Thicker valves enhance the jet intensity at the valve orifice and increase local flow velocity, while moderate thickness helps to reduce jet intensity and improve the uniformity of blood flow distribution. The second candidate leaflet model C exhibits lower peak levels in both wall pressure and wall shear stress, indicating its ability to effectively reduce local blood flow erosion. In conclusion, the determined leaflet thickness is 0.5 mm, and the second candidate leaflet model is selected as C.

[0285] In summary, this application demonstrates that increasing the radius of the leaflet free edge fillet significantly alleviates stress concentration and deformation accumulation, effectively suppressing local stress peaks. At a fillet radius of 3.5 mm, various strength indicators reach an optimal balance, achieving an overall improvement in structural mechanical performance without altering material parameters. With increasing fillet radius, the valve's geometric opening area significantly increases, and blood flow pattern analysis shows that smaller fillet radii tend to generate eccentric jets and localized turbulence, while a fillet radius of 3.5 mm results in the most uniform flow field distribution and significantly reduced helical flow. Furthermore, based on a fillet radius of 3.5 mm, increasing thickness enhances the overall valve stiffness and leads to a more uniform stress distribution. Moderately increasing thickness can reduce the maximum principal strain level and improve structural stability, but excessive thickness may affect the opening area and blood flow patency. Combining solid and fluid dynamics calculations, the medium-thickness model achieves a better balance between structural stability and opening efficiency.

[0286] This application also provides an optimized design system for the leaflets of a trilobed artificial biological aortic valve, which can be deployed, for example, in a computer device as a software tool or software module capable of processing data, performing data processing with the help of the hardware devices and / or the operating environment provided by the operating system in the computer device.

[0287] Please see Figure 49 The figure shows a block diagram of the optimization design system for the leaflets of a trilobed artificial biological aortic valve in one embodiment of this application. As shown in the figure, the optimization design system 1 includes a leaflet geometric model construction module 11, a first candidate model generation module 12, a first simulation evaluation module 13, a target fillet radius determination module 14, a second candidate model generation module 15, a second simulation evaluation module 16, and a target thickness determination module 17.

[0288] The system includes several modules: a leaflet geometry model construction module 11 for constructing a leaflet geometry model of a tricuspid artificial aortic valve, wherein the leaflet geometry model uses the leaflet free edge radius and leaflet thickness as variable geometric parameters; a first candidate model generation module 12 for generating multiple first candidate leaflet models by assigning multiple different leaflet free edge radii to the leaflet geometry model; a first simulation evaluation module 13 for performing multi-index simulation evaluations on the multiple first candidate leaflet models, including at least two of the following: ostium performance evaluation, structural mechanics evaluation based on the dynamic opening and closing process of the leaflets under transvalvular pressure differential, and hemodynamic evaluation based on an aortic flow field model; a target free edge radius determination module 14 for determining the target leaflet free edge radius based on the multi-index simulation evaluation results of the multiple first candidate leaflet models; and a second candidate model generation module 15 for generating multiple second candidate leaflet models by assigning multiple different leaflet thicknesses to the leaflet geometry model with the target leaflet free edge radius. A second simulation evaluation module 16 is used to perform the multi-index simulation evaluations on the multiple second candidate leaflet models. The target thickness determination module 17 is used to determine the target leaflet thickness based on the multi-index simulation evaluation results of multiple second candidate leaflet models, so as to obtain the target leaflet free edge radius and target leaflet thickness for the design of the trilobal artificial biological aortic valve leaflet.

[0289] In one embodiment, the optimization design system 1 includes a leaflet geometry model construction module 11, a first candidate model generation module 12, a first simulation evaluation module 13, a target fillet radius determination module 14, a second candidate model generation module 15, a second simulation evaluation module 16, and a target thickness determination module 17, which respectively coordinate and execute the optimization design method disclosed in any of the foregoing embodiments of this application according to the functions described above. Please refer to the [specific implementation details]. Figures 1 to 48 Any embodiment described herein, for example, the modules respectively implement steps S100 to S160 of the aforementioned method.

[0290] The module can also be implemented in software running on different types of processors. For example, an executable code module may include one or more physical or logical blocks of computer instructions organized as objects, programs, or functions. However, the executable files of the module do not necessarily have to be physically located together, but may include different commands stored in different locations, which, when logically connected together, encompass the module and implement the module's specified objectives.

[0291] Of course, the executable code module can be one or more instructions, and can even be distributed across several different code segments, different programs, and multiple storage devices. Similarly, computational data can be identified and represented within the module, and can be embodied in any suitable form and organized in any suitable data structure. The computational data can be collected as a single dataset, or can be distributed across different locations (including different storage devices), and can exist at least partially as electrical signals within a system or network. When the module or a portion thereof is implemented in software, the software portion is stored on one or more computer-readable media.

[0292] This application also provides a computer device; please refer to [link / reference]. Figure 50 The figure shows a schematic diagram of a computer device according to one embodiment of this application. As shown, the computer device 3 includes a storage device 31 and a processing device 32 connected to the storage device 31. Further, the computer device also includes a communication interface 33. The processing device 32 is used to call and execute the at least one program from the storage device 31 to implement the optimized design method of the tricuspid artificial aortic valve leaflets.

[0293] In some embodiments, the processing device 32 and the storage device 31 in the computer device 3 are interconnected at high speed via a system bus. The computer device may be a server, laptop computer, desktop computer, edge device, etc., and this application embodiment does not specifically limit it, nor does it limit the number of processors and memories in the computing device.

[0294] A bus can be a Peripheral Component Interconnect (PCI) bus or an Extended Industry Standard Architecture (EISA) bus, etc. Buses can be categorized as address buses, data buses, control buses, etc. For ease of representation, only one line is used in the diagram, but this does not imply that there is only one bus or one type of bus. A bus can include pathways for transmitting information between various components of a computing device (e.g., memory, processor, communication interfaces).

[0295] The processing device 32 may include any one or more processors such as a central processing unit (CPU), a graphics processing unit (GPU), a microprocessor (MP), or a digital signal processor (DSP). Alternatively, it may be an application-specific integrated circuit (ASIC), a discrete gate or transistor logic device, or a discrete hardware component, capable of implementing or executing the methods, steps, and logic block diagrams disclosed in the embodiments of this application.

[0296] Storage device 31 may include volatile memory, such as random access memory (RAM). Storage device 31 may also include non-volatile memory, such as read-only memory (ROM), random access memory (RAM), flash memory, programmable read-only memory (PROM), erasable programmable read-only memory (EPROM), electrically erasable programmable read-only memory (EEPROM), hard disk drive (HDD), or solid state drive (SSD). Storage device 31 is used to store executable program code, which is executed by processing device 32 upon receiving an execution instruction. Processing device 32 executes the executable program code to implement this application. Figure 1 The method steps in the embodiments.

[0297] In some embodiments, the communication interface 33 includes at least one interface unit, each interface unit being used to output a visual interface, receive human-computer interaction events generated according to the operation of a technician, etc. For example, the communication interface 33 includes, but is not limited to, serial interfaces such as HDMI interfaces or USB interfaces, or parallel interfaces, etc. In one embodiment, the communication interface 33 further includes a network communication unit, which is a device for data transmission using wired or wireless networks, examples of which include, but are not limited to, integrated circuits including network cards, local area network modules such as WiFi modules or Bluetooth modules, and wide area network modules such as mobile networks, etc.

[0298] This application also provides a computer-readable storage medium storing at least one program that, when executed by a computer processor, implements the optimized design method for the leaflets of the trilobed artificial biological aortic valve described in any of the above embodiments.

[0299] If the aforementioned functions are implemented as software functional units and sold or used as independent products, they can be stored in a computer-readable storage medium. Based on this understanding, the technical solution of this application, in essence, or the part that contributes to the prior art, or a part of the technical solution, can be embodied in the form of a software product. This computer software product is stored in a storage medium and includes several instructions to enable a computer device with the storage medium installed to execute all or part of the steps of the methods described in the various embodiments of this application.

[0300] Those skilled in the art will understand that implementing all or part of the processes in the methods described in the above embodiments can be accomplished by a computer program instructing related hardware. The computer program can be stored in a non-volatile computer-readable storage medium, and when executed, it can include the processes of the embodiments of the methods described above. Any references to memory, storage, databases, or other media used in the embodiments provided in this application can include non-volatile and / or volatile memory. Non-volatile memory may include read-only memory (ROM), programmable ROM (PROM), electrically programmable ROM (EPROM), electrically erasable programmable ROM (EEPROM), or flash memory, etc. Volatile memory may include random access memory (RAM) or external cache memory, etc. By way of illustration and not limitation, RAM is available in a variety of forms, such as static RAM (SRAM), dynamic RAM (DRAM), synchronous DRAM (SDRAM), dual data rate SDRAM (DDRSDRAM), enhanced SDRAM (ESDRAM), synchronous link DRAM (SLDRAM), RAMbus direct RAM (RDRAM), direct memory bus dynamic RAM (DRDRAM), and memory bus dynamic RAM (RDRAM), etc.

[0301] In the embodiments provided in this application, the computer-readable storage medium may include a read-only memory, random access memory, EEPROM, CD-ROM or other optical disc storage device, disk storage device or other magnetic storage device, flash memory, USB flash drive, portable hard drive, or any other medium capable of storing desired program code in the form of instructions or data structures and accessible by a computer. Additionally, any connection may be appropriately referred to as a computer-readable medium. For example, if instructions are transmitted from a website, server, or other remote source using coaxial cable, fiber optic cable, twisted pair, digital subscriber line (DSL), or wireless technologies such as infrared, radio, and microwave, then the coaxial cable, fiber optic cable, twisted pair, DSL, or wireless technologies such as infrared, radio, and microwave are included in the definition of the medium. However, it should be understood that computer-readable storage medium and data storage medium do not include connections, carrier waves, signals, or other transient media, but are intended for non-transient, tangible storage media. The disks and optical discs used in the application include compact optical discs (CDs), laser optical discs, optical discs, digital multifunction optical discs (DVDs), floppy disks, and Blu-ray discs, where disks typically copy data magnetically, while optical discs use lasers to copy data optically.

[0302] This application also provides a computer program product that, when run on a computer, causes the computer to execute the optimized design method for the leaflets of the trilobed artificial biological aortic valve of this application.

[0303] Those skilled in the art will understand that all or part of the processes in the above embodiments can be implemented by a computer program instructing related hardware. The computer program can be stored in a non-volatile computer-readable storage medium, and when executed, it can include the processes of the embodiments of the above methods. Any references to memory, databases, or other media used in the embodiments provided in this application can include at least one of non-volatile memory and volatile memory. Non-volatile memory can include read-only memory (ROM), magnetic tape, floppy disk, flash memory, optical memory, high-density embedded non-volatile memory, resistive random access memory (ReRAM), magnetic random access memory (MRAM), ferroelectric random access memory (FRAM), phase change memory (PCM), graphene memory, etc. Volatile memory can include random access memory (RAM) or external cache memory, etc. By way of illustration and not limitation, RAM can take many forms, such as Static Random Access Memory (SRAM) or Dynamic Random Access Memory (DRAM). The databases involved in the embodiments provided in this application may include at least one type of relational database and non-relational database. Non-relational databases may include, but are not limited to, blockchain-based distributed databases. The processors involved in the embodiments provided in this application may be general-purpose processors, central processing units, graphics processing units, digital signal processors, programmable logic devices, quantum computing-based data processing logic devices, artificial intelligence (AI) processors, etc., and are not limited to these.

[0304] In one or more exemplary aspects, the functions described in the computer program of the optimized design method for the tricuspid artificial aortic valve leaflets of this application can be implemented in hardware, software, firmware, or any combination thereof. When implemented in software, these functions can be stored or transmitted as one or more instructions or code onto a computer-readable medium. The steps of the methods or algorithms disclosed in this application can be embodied in processor-executable software modules, wherein the processor-executable software modules can reside on a tangible, non-transitory computer-readable storage medium. The tangible, non-transitory computer-readable storage medium can be any available medium accessible to a computer.

[0305] The flowcharts and block diagrams in the foregoing figures illustrate the architecture, functionality, and operation of possible implementations of systems, methods, and computer program products according to various embodiments of this application. Accordingly, each block in a flowchart or block diagram may represent a module, segment, or portion of code containing one or more executable instructions for implementing a specified logical function. It should also be noted that in some alternative implementations, the functions indicated in the blocks may occur in a different order than those indicated in the figures. For example, two consecutively indicated blocks may actually be executed substantially in parallel, and they may sometimes be executed in reverse order, depending on the functions involved. It should also be noted that each block in the block diagrams and / or flowcharts, and combinations of blocks in the block diagrams and / or flowcharts, can be implemented by a dedicated hardware-based system that performs the specified function or operation, or by a combination of dedicated hardware and computer instructions.

[0306] Those skilled in the art will recognize that the units and algorithm steps of the various examples described in conjunction with the embodiments disclosed herein can be implemented in electronic hardware, or a combination of computer software and electronic hardware. Whether these functions are implemented in hardware or software depends on the specific application and design constraints of the technical solution. Those skilled in the art can use different methods to implement the described functions for each specific application, but such implementation should not be considered beyond the scope of this application.

[0307] In summary, this application discloses an optimization design method, system, computer device, computer-readable storage medium, and computer program product for tricuspid artificial aortic valve leaflets. Using the leaflet free edge radius and leaflet thickness as two independent geometric design variables, the method first compares multiple candidate leaflet models with different free edge radii while keeping the leaflet thickness constant to determine the target leaflet free edge radius. Then, while keeping the target leaflet free edge radius unchanged, it compares multiple candidate leaflet models with different leaflet thicknesses to determine the target leaflet thickness. Compared to optimizing multiple geometric parameters simultaneously, this application employs a staged, single-variable optimization strategy, which effectively reduces the coupling effect between different design variables. This allows performance differences between different candidate models to be primarily attributed to changes in the corresponding design variables, thereby improving the reliability and interpretability of the parameter optimization results.

[0308] Furthermore, this application not only evaluates the geometric opening area of ​​the valve in the open state, but also combines the stress, strain, blood flow velocity, wall pressure, and wall shear stress in the aortic flow field during the dynamic opening and closing process of the valve leaflet under the action of transvalvular pressure difference to comprehensively evaluate different candidate models. This avoids selecting design parameters based solely on a single evaluation index, thereby taking into account valve opening performance, structural stability, and hemodynamic performance, and improving the rationality of the target leaflet free edge radius and target leaflet thickness.

[0309] This application obtains material parameters based on uniaxial tensile experiments of bovine pericardium and uses a hyperelastic constitutive model to characterize the valve leaflet material, enabling the valve leaflet material to more accurately reflect the large deformation nonlinear mechanical response of the artificial biological aortic valve during the cardiac cycle. Compared with using empirical material parameters or linear elastic material models, this application can improve the consistency between structural mechanics simulation analysis and actual working conditions, thereby improving the credibility of parameter optimization results.

[0310] This application, based on the comprehensive evaluation results of multiple candidate leaflet models, further screens candidate models that meet the preset opening performance conditions by combining structural mechanics evaluation results and hemodynamic evaluation results to determine the target leaflet free edge radius and target leaflet thickness. By establishing a unified evaluation process and parameter screening rules, the determination process of leaflet geometric parameters has a clear design basis, reducing the uncertainty brought about by empirical design and improving the standardization, repeatability, and engineering application value of the artificial biological aortic valve leaflet optimization design process.

[0311] The above embodiments are merely illustrative of the inventive essence and beneficial effects of this application, and are not intended to limit this application. Any person skilled in the art can modify or alter the above embodiments without departing from the principles and scope of this application. Therefore, all equivalent modifications or alterations achieved by those skilled in the art without departing from the spirit and technical concept disclosed in this application should still be covered by the claims of this application.

Claims

1. An optimized design method for the leaflets of a trilobate artificial biological aortic valve, characterized in that, The optimization design method includes the following steps: A leaflet geometric model of a trilobate artificial biological aortic valve is constructed, wherein the leaflet geometric model uses the radius of the free edge radius and the leaflet thickness as variable geometric parameters; Based on the aforementioned leaflet geometric model, multiple different leaflet free edge fillet radii are assigned to generate multiple first candidate leaflet models; Multiple candidate leaflet models are subjected to multi-index simulation evaluation, which includes at least two of the following: opening performance evaluation, structural mechanics evaluation based on the dynamic opening and closing process of the leaflet under transvalvular pressure difference, and hemodynamic evaluation based on the aortic flow field model. The radius of the free edge radius of the target leaflet is determined based on the multi-index simulation evaluation results of multiple first candidate leaflet models; Based on the leaflet geometric model with the radius of the free edge radius of the target leaflet, multiple different leaflet thicknesses are assigned to generate multiple second candidate leaflet models; The multi-index simulation evaluation was performed on multiple second candidate leaflet models; and Based on the multi-index simulation evaluation results of multiple second candidate leaflet models, the target leaflet thickness is determined to obtain the target leaflet free edge radius and target leaflet thickness for the design of the trilobal artificial biological aortic valve.

2. The optimization design method according to claim 1, characterized in that, The leaflet geometry model includes a basic leaflet unit, which includes a fixed edge, a ventral curved surface, and a free edge. The three leaflets are formed by circumferentially arranging the basic leaflet unit around the central axis of the trilobed artificial biological aortic valve.

3. The optimization design method according to claim 2, characterized in that, The geometric model of the basic leaflet unit is constructed by using two opposing boundary lines as the boundaries of the basic leaflet unit, and forming the free edge by connecting the two boundary lines with a circular arc, the radius of which is the fillet radius of the leaflet free edge.

4. The optimization design method according to claim 3, characterized in that, The geometric model of the basic leaflet unit further includes: rotating the arc forming the free edge about an axis to generate the ventral surface of the basic leaflet unit; wherein the axis is a line connecting the ends of the two boundary lines that are away from the arc.

5. The optimization design method according to claim 4, characterized in that, The arc rotates about the axis by an angle of 120°, so that the basic leaflet unit forms a composite surface with radial spherical curvature and circumferential arc curvature.

6. The optimization design method according to claim 1, characterized in that, Multiple first candidate leaflet models have the same leaflet thickness and the same leaflet geometric topology, and multiple second candidate leaflet models all have the target leaflet free edge fillet radius and the same leaflet geometric topology.

7. The optimization design method according to claim 1, characterized in that, The multi-index simulation evaluation includes the opening performance evaluation, the structural mechanics evaluation, and the hemodynamic evaluation.

8. The optimization design method according to claim 1 or 7, characterized in that, The opening performance evaluation includes determining the geometric opening area of ​​the first candidate leaflet model and the second candidate leaflet model in the open state, respectively. The geometric opening area is the valve opening area formed when the first candidate leaflet model and the second candidate leaflet model reach their maximum opening degree during the cardiac cycle.

9. The optimization design method according to claim 1 or 7, characterized in that, The structural mechanics evaluation includes determining the maximum principal stress, maximum principal strain, von Mises equivalent stress, and maximum shear stress of the first candidate leaflet model and the second candidate leaflet model during the dynamic opening and closing process.

10. The optimization design method according to claim 9, characterized in that, The structural mechanical evaluation includes obtaining the time-history peak values ​​and / or spatial average values ​​of the maximum principal stress, the maximum principal strain, the von Mises equivalent stress, and the maximum shear stress during the cardiac cycle.

11. The optimization design method according to claim 1 or 7, characterized in that, The hemodynamic evaluation includes determining the blood flow velocity, wall pressure, and wall shear stress corresponding to the first candidate leaflet model and the second candidate leaflet model after they are placed into the aortic flow field model.

12. The optimization design method according to claim 1, characterized in that, The various free edge radii of the leaflets are selected from multiple values ​​within the range of 1 mm to 3.5 mm.

13. The optimization design method according to claim 1, characterized in that, The various leaflet thicknesses are selected from multiple values ​​within the range of 0.4 mm to 0.6 mm.

14. The optimization design method according to claim 1, characterized in that, The structural mechanics evaluation includes: applying transvalvular pressure differential loads during the cardiac cycle to the first candidate leaflet model and the second candidate leaflet model, and performing finite element dynamic opening and closing analysis on the first candidate leaflet model and the second candidate leaflet model.

15. The optimization design method according to claim 1, characterized in that, The material parameters of the leaflet geometric model are obtained based on uniaxial tensile tests of bovine pericardium material, and the leaflet material in the first candidate leaflet model and the second candidate leaflet model is characterized by a hyperelastic constitutive model.

16. The optimization design method according to claim 15, characterized in that, The hyperelastic constitutive model is a second-order Ogden constitutive model.

17. The optimization design method according to claim 1, characterized in that, The hemodynamic evaluation includes: reconstructing the aortic flow field model based on aortic medical imaging data, and placing the first candidate leaflet model and the second candidate leaflet model into the aortic flow field model for computational fluid dynamics analysis.

18. The optimization design method according to claim 17, characterized in that, The aortic flow field model includes a left ventricular outflow tract inlet and multiple aortic outlets. The computational fluid dynamics analysis includes using the left ventricular outflow tract inlet as a velocity inlet and the multiple aortic outlets as pressure outlets.

19. The optimization design method according to claim 17 or 18, characterized in that, In the computational fluid dynamics analysis, blood is set as a uniform, incompressible, isotropic Newtonian fluid, and the aortic wall is set as a rigid, non-slip wall.

20. The optimization design method according to claim 7, characterized in that, Determining the radius of the free edge of the target leaflet includes: selecting first candidate leaflet models whose opening performance evaluation results meet the first preset opening conditions based on the multi-index simulation evaluation results of multiple first candidate leaflet models, and determining the radius of the free edge of the target leaflet among the selected first candidate leaflet models based on the structural mechanics evaluation results and hemodynamics evaluation results.

21. The optimization design method according to claim 20, characterized in that, Among the selected first candidate leaflet models, the leaflet free edge fillet radius corresponding to the first candidate leaflet model that satisfies the first preset structural conditions and the first preset flow field conditions is determined as the target leaflet free edge fillet radius; wherein, the first preset structural conditions include: the stress and / or strain level characterized by the structural mechanical evaluation result of the first candidate leaflet model determined as the target leaflet free edge fillet radius is lower than the stress and / or strain level of at least one first candidate leaflet model with a smaller leaflet free edge fillet radius.

22. The optimization design method according to claim 7, characterized in that, Determining the target leaflet thickness includes: selecting second candidate leaflet models whose opening performance evaluation results meet the second preset opening conditions based on the multi-index simulation evaluation results of multiple second candidate leaflet models, and determining the target leaflet thickness among the selected second candidate leaflet models based on the structural mechanics evaluation results and hemodynamics evaluation results.

23. The optimization design method according to claim 22, characterized in that, Among the selected second candidate leaflet models, the leaflet thickness corresponding to the second candidate leaflet model that satisfies the second preset structural conditions and the second preset flow field conditions is determined as the target leaflet thickness; wherein, the second preset structural conditions include: the stress and / or strain level characterized by the structural mechanical evaluation results of the second candidate leaflet model determined as the target leaflet thickness is lower than the stress and / or strain level of at least one second candidate leaflet model with a smaller leaflet thickness; the second preset flow field conditions include: the wall shear stress characterized by the hemodynamic evaluation results of the second candidate leaflet model determined as the target leaflet thickness satisfies the preset wall shear stress condition.

24. The optimization design method according to claim 23, characterized in that, The second preset flow field condition includes: the wall shear stress corresponding to the second candidate leaflet model that is determined to be the target leaflet thickness is lower than the wall shear stress corresponding to at least one second candidate leaflet model with a leaflet thickness less than the target leaflet thickness and at least one second candidate leaflet model with a leaflet thickness greater than the target leaflet thickness.

25. An optimized design system for the leaflets of a trilobate artificial biological aortic valve, characterized in that, include: The leaflet geometry model construction module is used to construct the leaflet geometry model of the trilobal artificial biological aortic valve. The leaflet geometry model uses the leaflet free edge radius and leaflet thickness as variable geometric parameters. The first candidate model generation module is used to generate multiple first candidate leaflet models by assigning multiple different free edge fillet radii to the leaflet geometric model. The first simulation evaluation module is used to perform multi-index simulation evaluation on multiple first candidate leaflet models. The multi-index simulation evaluation includes at least two of the following: opening performance evaluation, structural mechanics evaluation based on the dynamic opening and closing process of the leaflet under the action of transvalvular pressure difference, and hemodynamic evaluation based on the aortic flow field model. The target fillet radius determination module is used to determine the fillet radius of the free edge of the target leaflet based on the multi-index simulation evaluation results of multiple first candidate leaflet models; The second candidate model generation module is used to generate multiple second candidate leaflet models by assigning multiple different leaflet thicknesses to a leaflet geometric model with the radius of the free edge of the target leaflet. The second simulation evaluation module is used to perform the multi-index simulation evaluation on multiple second candidate leaflet models. The target thickness determination module is used to determine the target leaflet thickness based on the multi-index simulation evaluation results of multiple second candidate leaflet models, so as to obtain the target leaflet free edge radius and target leaflet thickness for the design of the trilobal artificial biological aortic valve leaflet.

26. A computer device, characterized in that, include: Storage device for storing at least one program; A processing device, connected to the storage device, is used to call and execute the at least one program from the storage device to implement the optimized design method of the trilobed artificial biological aortic valve leaflets as described in any one of claims 1 to 24.

27. A computer-readable storage medium, characterized in that, The computer-readable storage medium stores at least one program, which, when executed by a processor, implements the optimized design method for the leaflets of the trilobed artificial biological aortic valve as described in any one of claims 1 to 24.

28. A computer program product, characterized in that, The computer program product includes a computer program or instructions that, when executed by a processor, implement the optimized design method for the leaflets of the trilobed artificial biological aortic valve as described in any one of claims 1 to 24.

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