A biomechanical analysis system for talus prosthesis stability and stress distribution

Abstract

The present application relates to the technical field of prosthesis mechanics analysis, in particular to a kind of talus prosthesis stability and stress distribution biomechanics analysis system, by three-dimensional modeling module, finite element analysis module, stability analysis module and report generation module composition;Three-dimensional modeling module processes CT data by reverse engineering technology;The three-dimensional model of talus prosthesis is constructed;Finite element analysis module is used to simulate the stress condition of talus prosthesis in different positions and different thickness cartilage cover;Stability analysis module is used to analyze the stress distribution of talus prosthesis in different positions by finite element calculation after ligament is successively cut off;Report generation module is used to generate stress distribution nephogram, and the stress distribution of talus prosthesis surface is shown.The present application establishes the model of talus prosthesis, predicts stress distribution and identifies potential failure point or excessive wear based on finite element analysis, to guide design improvement and material selection.

Description

Technical Field

[0001] This invention relates to the field of prosthesis biomechanical analysis technology, specifically to a biomechanical analysis system for the stability and stress distribution of a talus prosthesis. Background Technology

[0002] The talus is an important bone in the human foot, connecting the lower limb and the foot. It has an irregular shape, is covered by five articular surfaces, and involves multiple movable joints, serving as a biomechanical transition point between the lower limb and the foot. Numerous ligaments and tendons pass around the talus to maintain its stability and motor function. Due to its unique anatomical and biomechanical properties, the talus is prone to ischemic necrosis under conditions such as trauma and autoimmune diseases. In severe cases, this can lead to talus collapse, significantly affecting a patient's standing and walking abilities.

[0003] In current techniques, arthrodesis is commonly used to treat talus collapse necrosis. While this method can relieve pain, it sacrifices ankle function, limiting the patient's range of activities in daily life, particularly noticeable in activities like climbing stairs and running. Furthermore, arthrodesis can lead to compensatory degeneration of surrounding joints, as well as muscle atrophy and osteoporosis due to prolonged immobilization.

[0004] In summary, how to solve the problem of limiting the range of motion of the joint when using arthroplasty to treat talus collapse necrosis has become an urgent problem to be solved in this field. Therefore, it is necessary to propose a biomechanical analysis system for the stability and stress distribution of talus prosthesis. Summary of the Invention

[0005] To address the aforementioned issues, this invention provides a biomechanical analysis system for the stability and stress distribution of talus prostheses. By establishing a model of the talus prosthesis, the system predicts stress distribution and identifies potential failure points or excessive wear based on finite element analysis, thereby guiding design improvements and material selection.

[0006] To achieve the above objectives, the technical solution of the present invention is as follows: a biomechanical analysis system for the stability and stress distribution of a talus prosthesis, comprising a three-dimensional modeling module, a finite element analysis module, a stability analysis module, and a report generation module.

[0007] The 3D modeling module is used to acquire CT data of the ankle joint using CT scans, process the CT data through reverse engineering techniques, construct a 3D model of the talus prosthesis, and design a geometric model of the talus prosthesis based on the geometry of the talus, tibia, and calcaneus.

[0008] The finite element analysis module is used to import 3D models into finite element analysis software for mesh generation and to apply material properties; it sets boundary conditions and loads based on human physiology and biomechanics principles to simulate the stress on the talus prosthesis at different locations and with different thicknesses of cartilage overlay.

[0009] The stability analysis module is used to analyze the stress distribution at different locations after the talus prosthesis is sequentially cut with different ligament combinations using finite element analysis; and to evaluate the overall stability of the talus prosthesis by combining the stress distribution results with the geometry of the talus prosthesis.

[0010] The report generation module generates stress distribution cloud maps to show the stress distribution on the surface of the talus prosthesis; and generates data analysis reports based on the results of finite element analysis.

[0011] Furthermore, in the 3D modeling module, reverse engineering technology processing includes the following steps:

[0012] S1, Data Acquisition and Processing: Import the raw DICOM format data from the CT scan into Mimics software, identify different parts of the bone and cartilage by setting different gray values ​​to obtain the initial model; export the talus, tibia and calcaneus as STL format model data files.

[0013] S2, Surface Treatment and Optimization: Use the "Relax", "Remove Spikes", and "Remove Filters" commands to fine-tune the initial model surface while preserving the model's features; import the STL file into Geomagic software, align the mesh with its exterior, and save it as an STP file.

[0014] S3, Model Import and Assembly: Import the STP file into SolidWorks for assembly and error checking; use the diagnostic function to check for distorted or missing surfaces, and save each bone part as an SLDPRT file after checking; use the assembly method to place each bone in the corresponding position; after assembly, use the interference check to identify and record overlapping parts.

[0015] S4, Create a 3D model: Use SolidWorks to create a ligament model, which includes the lateral, medial, talonavicular, and sinus tarsal ligaments; add the skin to the completed ankle model, and save the model as an X_T file.

[0016] Furthermore, in the finite element analysis module, the material properties are set as titanium alloy, with an elastic modulus of 105000 MPa and a density of 4.5 g / cm³. 3 The Poisson's ratio is 0.37.

[0017] Furthermore, in the finite element analysis module, the boundary conditions are set to insert fixed constraints at the upper ends of the tibia and fibula, and the load is set to apply a force of 500N to the foot.

[0018] Furthermore, in the finite element analysis module, the different positions of the talus prosthesis include neutral position, 5° dorsiflexion, 10° dorsiflexion, 5° plantarflexion, 10° plantarflexion, 20° plantarflexion, and 30° plantarflexion.

[0019] Furthermore, in the finite element analysis module, the talus prosthesis cartilage covering layer includes thicknesses of +0.5mm, +1.0mm, +1.5mm, and +2.0mm.

[0020] Furthermore, in the stability analysis module, after sequentially cutting the lateral-medial-tarsal sinus-talonavicular ligament, the range of motion of the talus prosthesis in ankle joint dorsiflexion 5°, dorsiflexion 10°, neutral position, plantarflexion 5°, plantarflexion 10°, plantarflexion 20°, and plantarflexion 30° was tested to measure the changes in anteroposterior displacement, inversion / exversion angle, internal / external rotation angle, and dorsal displacement distance of the talus head.

[0021] Furthermore, in the stability analysis module, after sequentially cutting the medial-tarsal sinus-talonavicular-lateral ligaments, the range of motion of the talus prosthesis was tested in ankle joint dorsiflexion 5°, dorsiflexion 10°, neutral position, plantarflexion 5°, plantarflexion 10°, plantarflexion 20°, and plantarflexion 30° to measure changes in anteroposterior displacement, inversion / exversion angle, internal / external rotation angle, and dorsal displacement distance of the talus head.

[0022] Furthermore, in the stability analysis module, after sequentially severing the sinus tarsi, talonavicular, lateral, and medial ligaments, the range of motion of the talus prosthesis was tested in ankle joint dorsiflexion 5°, dorsiflexion 10°, neutral position, plantarflexion 5°, plantarflexion 10°, plantarflexion 20°, and plantarflexion 30° to measure changes in anteroposterior displacement, inversion / exversion angle, internal / external rotation angle, and dorsal displacement distance of the talus head.

[0023] Furthermore, in the stability analysis module, after sequentially cutting the talonavicular-lateral-medial-tarsal sinus ligaments, the range of motion of the talus prosthesis was tested in ankle joint dorsiflexion 5°, dorsiflexion 10°, neutral position, plantarflexion 5°, plantarflexion 10°, plantarflexion 20°, and plantarflexion 30° to measure changes in anteroposterior displacement, inversion / exversion angle, internal / external rotation angle, and dorsal displacement distance of the talus head.

[0024] The above approach has the following beneficial effects:

[0025] 1. This invention applies finite element analysis to post-placement testing of the talus prosthesis. The process begins with creating a detailed and accurate model of the ankle joint, including the implanted prosthesis and surrounding bone and tissue structures. Using high-resolution imaging data and advanced modeling software, precise representations of anatomical features can be generated. Insights gained from finite element analysis simulations can inform clinical practice. Surgeons can use this information to make more informed decisions regarding prosthesis selection, placement techniques, and postoperative care. Understanding the biomechanical interactions within the ankle joint can optimize surgical outcomes and promote patient recovery.

[0026] In summary, the use of finite element analysis (FEM) in the evaluation of talus prostheses is a crucial step in improving the design and performance of these implants. By providing a detailed understanding of the prosthesis's mechanical behavior and its interaction with surrounding structures, FEM helps ensure patients receive durable, functional, and reliable solutions for severe talus injuries. It allows for a detailed assessment of the prosthesis's structural integrity and its interaction with surrounding bone and tissues. Secondly, FEM helps predict stress distribution and identify potential failure points or excessive wear, guiding design improvements and material selection. It provides insights into the prosthesis's performance under different loading conditions, which is essential for ensuring its durability and effective functional restoration.

[0027] 2. This invention utilizes various finite element analysis software to process the talus prosthesis. Based on the raw DICOM format data from CT scans, it identifies and constructs high-precision initial models of the talus, tibia, and calcaneus. Optimization processing allows for fine-tuning of the initial model surface, removing unnecessary details and noise while preserving key model features; this simplifies model complexity, improves computational efficiency, and maintains the model's biological realism. Assembly and error checking using SolidWorks identify and record overlapping components after assembly, preventing biomechanical analysis errors caused by improper contact between parts. Incorporating soft tissues such as ligaments and skin into the 3D model more realistically simulates the actual state of the human ankle joint; this facilitates a more comprehensive assessment of the talus prosthesis's stability and stress distribution, improving the accuracy and reliability of the analysis.

[0028] 3. This invention analyzes the stress distribution of the talus prosthesis under cartilage overlays of different thicknesses (+0.5mm, +1.0mm, +1.5mm, and +2.0mm) to evaluate the impact of the cartilage overlay on prosthesis stability and stress distribution. This helps optimize the design of the cartilage overlay, improving the biocompatibility and lifespan of the prosthesis. By analyzing the stress distribution of the talus prosthesis in different positions (neutral, 5° dorsiflexion, 10° dorsiflexion, 5° plantarflexion, 10° plantarflexion, 20° plantarflexion, and 30° plantarflexion), a more comprehensive understanding of the prosthesis's mechanical properties and stability can be achieved. A comprehensive evaluation of the prosthesis's performance in various common physiological activities is of great significance for understanding the stability and stress distribution of the prosthesis under different movement patterns.

[0029] 4. This invention comprehensively evaluates the stability of the talus prosthesis by severing different combinations of lateral, medial, sinus tarsi, and talonavicular ligaments, and testing the displacement and angular changes of the talus prosthesis under different ranges of ankle joint movement. Comparing the effects of different ligament severance sequences on the stability of the talus prosthesis can provide directions for improvement in prosthesis design.

[0030] Additional aspects and advantages of the invention will be set forth in part in the description which follows, and in part will be obvious from the description, or may be learned by practice of the invention. Attached Figure Description

[0031] Figure 1 This is a structural block diagram of the biomechanical analysis system for the stability and stress distribution of the talus prosthesis in an embodiment of the present invention.

[0032] Figure 2 This is a schematic diagram of the reverse engineering process in an embodiment of the present invention. Detailed Implementation

[0033] The following detailed description illustrates the specific implementation method:

[0034] Example: As attached Figures 1-2 As shown: A biomechanical analysis system for the stability and stress distribution of a talus prosthesis, mainly composed of a three-dimensional modeling module, a finite element analysis module, a stability analysis module, and a report generation module.

[0035] The 3D modeling module is used to acquire CT data of the ankle joint using CT scans, process the CT data through reverse engineering techniques, construct a 3D model of the talus prosthesis, and design a geometric model of the talus prosthesis based on the geometry of the talus, tibia, and calcaneus.

[0036] The reverse engineering process includes the following steps:

[0037] S1, Data Acquisition and Processing: Import the raw DICOM format data from the CT scan into Mimics software, identify different parts of the bone and cartilage by setting different gray values ​​to obtain the initial model; export the talus, tibia and calcaneus as STL format model data files.

[0038] S2, Surface Treatment and Optimization: Use the "Relax", "Remove Spikes", and "Remove Filters" commands to fine-tune the initial model surface while preserving the model's features; import the STL file into Geomagic software, align the mesh with its exterior, and save it as an STP file.

[0039] S3, Model Import and Assembly: Import the STP file into SolidWorks for assembly and error checking; use the diagnostic function to check for distorted or missing surfaces, and save each bone part as an SLDPRT file after checking; use the assembly method to place each bone in the corresponding position; after assembly, use the interference check to identify and record overlapping parts.

[0040] S4, Create a 3D model: Use SolidWorks to create a ligament model, which includes the lateral, medial, talonavicular, and sinus tarsal ligaments; add the skin to the completed ankle model, and save the model as an X_T file.

[0041] The finite element analysis module is used to import the 3D model into the finite element analysis software for mesh generation and to apply material properties of titanium alloy, setting its elastic modulus to 105000 MPa and density to 4.5 g / cm³. 3 The Poisson's ratio is 0.37. Based on the principles of human physiology and biomechanics, the boundary conditions are set as fixed constraints inserted at the upper ends of the tibia and fibula, and the load is set as a force of 500N applied to the foot. The stress conditions of the talus prosthesis in different positions (neutral, 5° dorsiflexion, 10° dorsiflexion, 5° plantarflexion, 10° plantarflexion, 20° plantarflexion, and 30° plantarflexion) and the cartilage covering layers with different thicknesses of +0.5mm, +1.0mm, +1.5mm, and +2.0mm are simulated.

[0042] The stability analysis module is used to analyze the stress distribution at different locations after the talus prosthesis is sequentially cut with different ligament combinations using finite element analysis; and to evaluate the overall stability of the talus prosthesis by combining the stress distribution results with the geometry of the talus prosthesis.

[0043] The different ligament combinations involved sequentially cutting the lateral-medial-tarsal sinus-talonavicular ligament, the medial-tarsal sinus-talonavicular-lateral ligament, the tarsal sinus-talonavicular-lateral-medial ligament, and the talonavicular-lateral-medial-tarsal sinus ligament. The range of motion of the talus prosthesis was tested in ankle joint dorsiflexion 5°, dorsiflexion 10°, neutral position, plantarflexion 5°, plantarflexion 10°, plantarflexion 20°, and plantarflexion 30°. The changes in anteroposterior displacement, inversion / valgus angle, internal / external rotation angle, and dorsal displacement distance of the talus head were also measured.

[0044] The report generation module generates stress distribution cloud maps to show the stress distribution on the surface of the talus prosthesis; and generates data analysis reports based on the results of finite element analysis.

[0045] Experiment and demonstration

[0046] 1. Method

[0047] Data source

[0048] Adult 3D ankle CT scan files were obtained from the hospital through official channels. Before use, confirm that the 3D CT data is usable and sufficiently clear.

[0049] 2. Basic steps for establishing an ankle joint model:

[0050] S1, Data Acquisition and Processing: Import the raw DICOM format data from the CT scan into Mimics software, identify different parts of the bone and cartilage by setting different gray values ​​to obtain the initial model; export the talus, tibia and calcaneus as STL format model data files.

[0051] S2, Surface Treatment and Optimization: Use the "Relax", "Remove Spikes", and "Remove Filters" commands to fine-tune the initial model surface while preserving the model's features; import the STL file into Geomagic software, align the mesh with its exterior, and save it as an STP file.

[0052] S3, Model Import and Assembly: Import the STP file into SolidWorks for assembly and error checking; use the diagnostic function to check for distorted or missing surfaces, and save each bone part as an SLDPRT file after checking; use the assembly method to place each bone in the corresponding position; after assembly, use the interference check to identify and record overlapping parts.

[0053] S4. Create a 3D model: Use SolidWorks to create a ligament model, which includes the lateral, medial, talonavicular, and sinus tarsal ligaments; add the skin to the completed ankle joint model, save the model as an X_T file, and import it into AnsysWorkbench for analysis settings.

[0054] After model creation, the mechanical properties of bone, prosthesis, skin, ligaments, and cartilage were incorporated into the model. Different ankle joint positions were simulated under various ligament resection conditions, including neutral, 5° dorsiflexion, 10° dorsiflexion, 5° plantarflexion, 10° plantarflexion, 20° plantarflexion, and 30° plantarflexion. Surrounding ligaments, including the lateral, medial, sinus tarsi, and talonavicular ligaments, were sequentially resected. Cartilage overlays of varying thicknesses (+0.5mm, +1.0mm, +1.5mm, and +2.0mm) were then added.

[0055] 3. Finite element analysis:

[0056] Open Ansys Workbench and create a static analysis module. Import the STP file into the static module. Use tetrahedral meshing and adjust the amounts according to the model characteristics. Fine-tune the mesh empirically to improve quality and data accuracy. Apply the material property of titanium alloy. In the static structure module, right-click to insert fixed constraints at the upper ends of the tibia and fibula, applying a force of 500N to the foot. Insert the results of maximum equivalent stress and deformation. Select the appropriate model in the parameter settings to extract the desired results.

[0057] 4. Discussion of Results

[0058] 4.1 Taurus prosthesis +0mm

[0059] Overall displacement of the talus prosthesis varied across different positions (Table 1). In the neutral position, the displacement ranged from 1.2474 mm to 1.7748 mm. Displacement increased during dorsiflexion, reaching 2.0935 mm at 5° and 2.125 mm at 10°. During plantarflexion, displacement decreased with increasing angle: 1.524 mm at 5°, 1.549 mm at 10°, 0.97307 mm at 20°, and 0.8736 mm at 30°. These results indicate that displacement was most significant during dorsiflexion and minimal during high plantarflexion.

[0060] The X-axis displacement varies with position. In the neutral position, it is -0.12642 mm, increasing to -0.23578 mm at 5° dorsiflexion and to -0.19279 mm at 10° dorsiflexion. In plantar flexion, the X-axis displacement changes from -0.052018 mm at 5° to -0.074974 mm at 10°, then to a positive value of 0.039485 mm at 20° and 0.042872 mm at 30°. This change from negative to positive values ​​indicates that the prosthesis position changes with increasing plantar flexion.

[0061] The Y-axis displacement is consistently negative, indicating downward movement. It is -1.2321 mm in neutral position, increasing to -1.4268 mm and -1.4545 mm at 5° and 10° dorsiflexion, respectively. In plantar flexion, the displacement is -1.065 mm at 5°, -1.0617 mm at 10°, decreasing to -0.65525 mm at 20°, and -0.57521 mm at 30°. This reflects the effects of gravity and flexion angle on the prosthesis.

[0062] The Z-axis displacement is positive at all positions, indicating upward movement. In the neutral position, it is 0.54856 mm, increasing to 0.6273 mm (5°) and 0.64081 mm (10°) during dorsiflexion. In plantarflexion, the displacement is 0.47063 mm at 5°, 0.481 mm at 10°, 0.3095 mm at 20°, and 0.27463 mm at 30°. This indicates that the Z-axis displacement decreases with increasing plantarflexion angle.

[0063] The contact area is largest in the neutral position, at 480.97 mm² (Table 1). It decreases slightly during dorsal extension, to 480.11 mm². 2 (5°) and 479.86mm 2 (10°). During plantar flexion, the contact area remains at 480.94 mm around the circumference. 2 (5°) and 480.93mm 2 (10°), but significantly reduced to 474.04 mm² (20°) and 445.35 mm² (30°). 2These changes indicate that contact is stable at lower buckling angles and more variable at larger angles.

[0064] Significant differences in contact stress were observed at different locations (Table 1). In the neutral position, the maximum contact stress was 6.61 MPa, and the minimum was 0.28073 MPa. During dorsiflexion, the stress decreased, reaching a maximum of 2.7272 MPa at 5° and 3.0024 MPa at 10°. During plantar flexion, the maximum stress was 2.0474 MPa and 2.0242 MPa at 5° and 10°, respectively, but increased with increasing angle, reaching 6.9746 MPa at 20° and 6.2178 MPa at 30°. The minimum stress was 0.18389 MPa. These results indicate the influence of buckling angle on stress distribution.

[0065] 4.2 Taurus prosthesis +0.5mm

[0066] The total displacement of the talus prosthesis with an eccentricity of 0.5 mm varied significantly across different positions (Table 2). In the neutral position, the displacement was 2.2572 mm. During dorsiflexion at 5° and 10°, the displacements were 2.1077 mm and 1.9777 mm, respectively. For plantar flexion at 5°, 10°, 20°, and 30°, the displacements were 1.5117 mm, 1.15 mm, 2.6925 mm, and 0.80259 mm, respectively. This indicates that the displacement was most pronounced during dorsiflexion and 20° of plantar flexion, and minimal during high plantar flexion.

[0067] The X-axis displacement varies significantly with position (Table 2). In the neutral position, the displacement is -0.05795 mm. During 5° dorsiflexion, this value increases to -0.27465 mm, and decreases to -0.14123 mm at 10°. During plantar flexion, the displacement changes from -0.054101 mm at 5° to 0.027218 mm at 10°, 0.14158 mm at 20°, and 0.20806 mm at 30°. This change from negative to positive values ​​with increasing plantar flexion angle indicates lateral movement of the prosthesis.

[0068] The Y-axis displacement shows consistently negative values, indicating downward movement (Table 2). In the neutral position, the displacement is -0.7499 mm. During dorsiflexion at 5° and 10°, it increases to -1.4337 mm and -1.3591 mm, respectively. In plantar flexion, the displacement is -1.0551 mm at 5°, -0.78776 mm at 10°, -2.0194 mm at 20°, and -0.50057 mm at 30°. These results reflect the effects of gravity and buckling angle on the prosthesis.

[0069] The Z-axis displacement shows a positive trend across all positions, indicating upward movement (Table 2). In the neutral position, the displacement is 0.33928 mm. At 5° and 10° of dorsiflexion, the displacements are 0.62143 mm and 0.60524 mm, respectively. In plantar flexion, the displacements are 0.4675 mm at 5°, 0.37588 mm at 10°, 0.70613 mm at 20°, and 0.27276 mm at 30°. This indicates that the Z-axis displacement varies with the buckling angle and is influenced by the direction of movement.

[0070] The neutral position has the largest contact area, at 481.64 mm. 2 (Table 2). During extension, the thickness decreased slightly to 480.96 mm at 5° and 10° respectively. 2 and 472.14mm 2 During plantar flexion, the contact area remained relatively stable at 5° and 10°, at 468.47 mm². 2 and 469.44mm 2 However, when reduced to 20°, it is 474.04 mm. 2 At 30°, it is 445.24 mm. 2 These changes indicate that contact is stable at lower buckling angles and more variable at larger angles.

[0071] The contact stress experienced by the talus prosthesis varied at different positions (Table 2). In the neutral position, the maximum contact stress was 1.5667 MPa, and the minimum was 0.11415 MPa. During dorsiflexion, the stress increased, reaching a maximum of 2.6961 MPa at 5° and 2.7708 MPa at 10°, with corresponding minimums of 0.32898 MPa and 0.24591 MPa. During plantar flexion, the maximum stress was 2.0466 MPa at 5°, 1.459 MPa at 10°, 2.434 MPa at 20°, and 1.1884 MPa at 30°, with a minimum of 0.17031 MPa. These results indicate the influence of different flexion angles on stress distribution, with higher angles leading to increased stress.

[0072] 4.3 Taurus prosthesis +1mm

[0073] The overall displacement of the 1 mm offset talus prosthesis showed substantial differences across different positions (Table 3). In the neutral position, the displacement was 1.7882 mm. During dorsiflexion at 5° and 10°, the displacements were 2.0337 mm and 2.126 mm, respectively. For plantar flexion at 5°, 10°, 20°, and 30°, the displacements were 1.6886 mm, 1.2161 mm, 0.79363 mm, and 0.90878 mm, respectively. This indicates that the displacement was most significant during dorsiflexion and minimal during higher degrees of plantar flexion.

[0074] The X-axis displacement varies significantly with position (Table 3). In the neutral position, the displacement is -0.12261 mm. During 5° dorsiflexion, this value increases to -0.26078 mm, and decreases to -0.15449 mm at 10°. During plantar flexion, the displacement ranges from -0.10096 mm at 5° to 0.070816 mm at 10°, 0.042472 mm at 20°, and 0.055053 mm at 30°. This change from negative to positive values ​​with increasing plantar flexion angle indicates lateral movement of the prosthesis.

[0075] The Y-axis displacement shows a consistently negative value, indicating downward movement (Table 3). In the neutral position, the displacement is -1.2384 mm. During dorsiflexion at 5° and 10°, it increases to -1.4001 mm and -1.4614 mm, respectively. In plantar flexion, the displacement is -1.1789 mm at 5°, -0.83049 mm at 10°, -0.5495 mm at 20°, and -0.59572 mm at 30°. These results reflect the effects of gravity and buckling angle on the prosthesis.

[0076] The Z-axis displacement shows a positive trend across all positions, indicating upward movement (Table 3). In the neutral position, the displacement is 0.56644 mm. At 5° and 10° of dorsiflexion, the displacements are 0.59316 mm and 0.6469 mm, respectively. In plantar flexion, the displacements are 0.51125 mm at 5°, 0.39929 mm at 10°, 0.24884 mm at 20°, and 0.29685 mm at 30°. This indicates that the Z-axis displacement varies with the buckling angle and is influenced by the direction of movement.

[0077] The neutral position has the largest contact area, at 482.82 mm. 2 (Table 3). During extension, the thickness decreased slightly to 480.96 mm at 5° and 10°, respectively. 2 and 468.42mm 2 During plantar flexion, the contact area remained relatively stable at 5° and 10°, at 460.74 mm². 2 and 465.23mm 2 However, it significantly decreased to 474.04 mm at 20°. 2 At 30°, it is 445.47 mm. 2 These changes indicate that contact is stable at lower flexion angles, while greater variability occurs at higher angles. The talus prosthesis experiences different contact stresses at different locations (Table 3).

[0078] In the neutral position, the maximum contact stress was 3.6539 MPa, and the minimum was 0.17745 MPa. During extension, the stress increased, reaching a maximum of 2.595 MPa at 5° and 3.0038 MPa at 10°, with corresponding minimums of 0.26396 MPa and 0.23529 MPa, respectively. During plantar flexion, the maximum stress was 2.2812 MPa at 5°, 1.5661 MPa at 10°, 0.70867 MPa at 20°, and 1.371 MPa at 30°, with a minimum of 0.24633 MPa. These results reveal the influence of different buckling angles on stress distribution, with higher angles leading to increased stress.

[0079] 4.4 talus prosthesis + 1.5mm

[0080] The total displacement of the 1.5 mm offset talus prosthesis varied significantly across different positions (Table 4). In the neutral position, the displacement was 1.9302 mm. During dorsiflexion at 5° and 10°, the displacements were 2.1234 mm and 2.1022 mm, respectively. For plantar flexion at 5°, 10°, 20°, and 30°, the displacements were 1.671 mm, 1.4322 mm, 1.0149 mm, and 0.96717 mm, respectively. These findings suggest that the displacement was most pronounced during dorsiflexion and minimized at higher degrees of plantar flexion.

[0081] The X-axis displacement varies significantly with position (Table 4). In the neutral position, the displacement is -0.20596 mm. During dorsiflexion, it is -0.25359 mm at 5° and -0.19579 mm at 10°. During plantar flexion, the displacement changes from -0.09551 mm at 5° to -0.051167 mm at 10°, 0.068006 mm at 20°, and 0.043149 mm at 30°. This change from negative to positive values ​​with increasing plantar flexion angle indicates lateral movement of the prosthesis.

[0082] The Y-axis displacement shows a consistently negative value, indicating downward movement (Table 4). In the neutral position, the displacement is -1.326 mm. During dorsiflexion at 5° and 10°, it increases to -1.4455 mm and -1.4409 mm, respectively. In plantar flexion, the displacement is -1.1692 mm at 5°, -0.98634 mm at 10°, -0.68009 mm at 20°, and -0.63279 mm at 30°. These results reflect the effects of gravity and buckling angle on the prosthesis.

[0083] The Z-axis displacement shows a positive trend across all positions, indicating upward movement (Table 4). In the neutral position, the displacement is 0.58021 mm. At 5° and 10° of dorsiflexion, the displacements are 0.6327 mm and 0.62447 mm, respectively. In plantar flexion, the displacements are 0.50402 mm at 5°, 0.44701 mm at 10°, 0.33733 mm at 20°, and 0.31373 mm at 30°. This indicates that the Z-axis displacement varies with the buckling angle and is influenced by the direction of movement.

[0084] The neutral position has the largest contact area, at 481.64 mm. 2 (Table 4). During extension, the thickness decreased slightly to 480.97 mm at 5° and 10°, respectively. 2 and 470.55mm 2 During plantar flexion, the contact area remained relatively stable at 5° and 10°, at 468.43 mm². 2 and 469.46mm 2 However, it decreased to 474.04 mm at 20° and 30° respectively. 2 and 456.99mm 2 These changes indicate that contact is stable at lower flexion angles and more variable at larger angles. The talus prosthesis experiences different contact stresses at different locations (Table 4).

[0085] In the neutral position, the maximum contact stress is 3.0293 MPa, and the minimum is 0.3016 MPa. During extension, the stress increases, reaching a maximum of 2.7719 MPa at 5° and 3.2242 MPa at 10°, with corresponding minimums of 0.29949 MPa and 0.27412 MPa, respectively. During plantar flexion, the maximum stress is 2.2429 MPa at 5°, 1.7929 MPa at 10°, and 1.0711 MPa at 30°, with a minimum of 0.17669 MPa. These results indicate the influence of different buckling angles on stress distribution, with larger angles leading to increased stress.

[0086] 4.5 talus prosthesis + 2mm

[0087] The overall displacement of the 2mm offset talus prosthesis showed substantial differences across different positions (Table 5). In the neutral position, the displacement was 1.925mm. During dorsiflexion at 5° and 10°, the displacements were 2.2326mm and 2.1046mm, respectively. For plantar flexion at 5°, 10°, 20°, and 30°, the displacements were 1.5177mm, 1.2174mm, 1.1053mm, and 0.97152mm, respectively. These findings suggest that the displacement was most pronounced during dorsiflexion and minimized at higher degrees of plantar flexion.

[0088] The X-axis displacement varies significantly with position (Table 5). In the neutral position, the displacement is -0.16778 mm. During dorsiflexion, it is -0.20529 mm at 5° and -0.15239 mm at 10°. In plantarflexion, the displacement is -0.047877 mm at 5°, 0.084627 mm at 10°, 0.066384 mm at 20°, and 0.05696 mm at 30°. This change from negative to positive values ​​with increasing plantarflexion angle indicates lateral movement of the prosthesis.

[0089] The Y-axis displacement shows a consistently negative value, indicating downward movement (Table 5). In the neutral position, the displacement is -1.3287 mm. During dorsiflexion at 5° and 10°, it increases to -1.5315 mm and -1.4535 mm, respectively. In plantar flexion, the displacement is -1.0599 mm at 5°, -0.83033 mm at 10°, -0.7425 mm at 20°, and -0.63912 mm at 30°. These results reflect the effects of gravity and flexion angle on the prosthesis.

[0090] The Z-axis displacement shows a positive trend across all positions, indicating upward movement (Table 5). In the neutral position, the displacement is 0.58314 mm. At 5° and 10° of dorsiflexion, the displacements are 0.67009 mm and 0.63508 mm, respectively. In plantar flexion, the displacements are 0.47002 mm at 5°, 0.40181 mm at 10°, 0.36675 mm at 20°, and 0.3147 mm at 30°. This indicates that the Z-axis displacement varies with the buckling angle and is influenced by the direction of movement.

[0091] The neutral position has the largest contact area, at 481.63 mm. 2 (Table 5). During extension, the thickness decreased slightly to 480.94 mm at 5° and 10°, respectively. 2 and 468.56mm 2 During plantar flexion, the contact area remained relatively stable at 5° and 10°, at 468.44 mm². 2 and 469.48mm 2 However, it significantly decreased to 474.84 mm at 20° and 30° respectively. 2 and 458.05mm 2 These changes indicate that contact is stable at lower flexion angles and more variable at larger angles. The talus prosthesis experiences different contact stresses at different locations (Table 5).

[0092] In the neutral position, the maximum contact stress was 2.5605 MPa, and the minimum was 0.32597 MPa. During dorsiflexion, the stress increased, reaching a maximum of 2.9225 MPa at 5° and 2.8961 MPa at 10°, with corresponding minimums of 0.34511 MPa and 0.33263 MPa, respectively. The maximum plantar flexion stress was 1.009 MPa at 5°, 1.5694 MPa at 10°, 1.3441 MPa at 20°, and 1.3744 MPa at 30°, with a minimum of 0.087747 MPa. These results reveal the influence of different buckling angles on stress distribution, with higher angles leading to increased stress.

[0093] Table 1. Summary of displacement and contact stress of talus prosthesis at +0 mm

[0094]

[0095] Table 2. Summary of displacement and contact stress of talus prosthesis +0.5mm

[0096]

[0097]

[0098] Table 3. Summary of talus prosthesis displacement and contact stress at +1mm.

[0099]

[0100] Table 4. Summary of displacement and contact stress of talus prosthesis +1.5mm

[0101]

[0102]

[0103] Table 5. Summary of talus prosthesis displacement and contact stress (+2mm)

[0104]

[0105] Obviously, the above embodiments are merely illustrative examples for clear explanation and are not intended to limit the implementation. Those skilled in the art will recognize that other variations or modifications can be made based on the above description. It is neither necessary nor possible to exhaustively list all possible implementations here. However, obvious variations or modifications derived therefrom are still within the scope of protection of this invention.

Claims

1. A biomechanical analysis system for talus prosthesis stability and stress distribution, comprising a three-dimensional modeling module, a finite element analysis module, a stability analysis module, and a report generation module; characterized in that: the three-dimensional modeling module is used to obtain CT data of the ankle joint using CT scanning, process the CT data through reverse engineering technology; construct a three-dimensional model of the talus prosthesis; design a geometric model of the talus prosthesis according to the geometric shapes of the talus, tibia and calcaneus; the finite element analysis module is used to import the three-dimensional model into finite element analysis software for meshing and applying material properties; boundary conditions and loads are set according to the principles of human physiology and biomechanics to simulate the stress conditions of the talus prosthesis in different positions and different thickness of cartilage covering layers; the stability analysis module is used to analyze the stress distribution of the talus prosthesis in different positions after cutting different ligament combinations in turn through finite element calculation; combined with the stress distribution results and the geometric shape of the talus prosthesis, the overall stability of the talus prosthesis is evaluated; the report generation module is used to generate a stress distribution cloud chart to show the stress distribution on the surface of the talus prosthesis; and generate a data analysis report according to the results of the finite element analysis calculation.

2. The system for biomechanical analysis of talus prosthesis stability and stress distribution according to claim 1, characterized in that, In the three-dimensional modeling module, the reverse engineering technology processing includes the following steps: S1, data acquisition and processing: import the DICOM format original data of CT scanning into Mimics software, identify different parts of bone and cartilage by setting different gray values to obtain an initial model; export the talus, tibia and calcaneus as an STL format model data file; S2, surface treatment and optimization: use "relaxation", "remove spikes" and "remove filter" commands to fine-tune the surface of the initial model, while preserving the features of the model; import the STL file into Geomagic software to align the mesh with its exterior, and save it as an STP file; S3, model import and assembly: import the STP file into SolidWorks for assembly and error checking; use the diagnostic function to check for twisted or missing surfaces, and after checking, save each bone part as an SLDPRT file; use the assembly method to place each bone in the corresponding position; after assembly, use interference checking to identify and record overlapping parts; S4, create a three-dimensional model: use SolidWorks to create a ligament model, which includes lateral, medial, talonavicular and tarsal sinus ligaments; add skin to the completed ankle joint model, and save the model as an X_T format.

3. The system for biomechanical analysis of talus prosthesis stability and stress distribution of claim 2, wherein, In the finite element analysis module, the material property is titanium alloy, and the elastic modulus thereof is set to 105000 MPa, the density thereof is set to 4.5 g / cm 3 , and the Poisson's ratio thereof is set to 0.

37.

4. The system for biomechanical analysis of talus prosthesis stability and stress distribution according to claim 3, characterized in that, In the finite element analysis module, the boundary conditions are inserted fixed constraints on the upper end of the tibia and fibula, and the load is applied to the foot with a force of 500N.

5. The system for biomechanical analysis of talus prosthesis stability and stress distribution according to claim 4, characterized in that, In the finite element analysis module, the different positions of the talus prosthesis include neutral position, 5° dorsiflexion, 10° dorsiflexion, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion.

6. The system for biomechanical analysis of talus prosthesis stability and stress distribution according to claim 5, characterized in that, In the finite element analysis module, the cartilage covering layer of the talus prosthesis includes thicknesses of +0.5mm, +1.0mm, +1.5mm and +2.0mm.

7. The system for biomechanical analysis of talus prosthesis stability and stress distribution according to claim 6, characterized in that, In the stability analysis module, after cutting off the lateral-medial-tarsal sinus-talus navicular ligament in turn, the change of the anterior-posterior displacement of the talus, the medial-lateral tilt angle, the medial-lateral rotation angle and the dorsal displacement distance of the talus head of the talus prosthesis in the range of motion of the ankle joint dorsiflexion 5°, dorsiflexion 10°, neutral position, plantar flexion 5°, plantar flexion 10°, plantar flexion 20° and plantar flexion 30° is tested.

8. The system for biomechanical analysis of talus prosthesis stability and stress distribution of claim 6, wherein, In the stability analysis module, after cutting off the medial-tarsal sinus-talus navicular-lateral ligament in turn, the change of the anterior-posterior displacement of the talus, the medial-lateral tilt angle, the medial-lateral rotation angle and the dorsal displacement distance of the talus head of the talus prosthesis in the range of motion of the ankle joint dorsiflexion 5°, dorsiflexion 10°, neutral position, plantar flexion 5°, plantar flexion 10°, plantar flexion 20° and plantar flexion 30° is tested.

9. The system for biomechanical analysis of talus prosthesis stability and stress distribution of claim 6, wherein, In the stability analysis module, after cutting off the tarsal sinus-talus navicular-lateral-medial ligament in turn, the change of the anterior-posterior displacement of the talus, the medial-lateral tilt angle, the medial-lateral rotation angle and the dorsal displacement distance of the talus head of the talus prosthesis in the range of motion of the ankle joint dorsiflexion 5°, dorsiflexion 10°, neutral position, plantar flexion 5°, plantar flexion 10°, plantar flexion 20° and plantar flexion 30° is tested.

10. The system for biomechanical analysis of talus prosthesis stability and stress distribution of claim 6, wherein, In the stability analysis module, after cutting off the tarsal sinus-talus navicular-lateral-medial ligament in turn, the change of the anterior-posterior displacement of the talus, the medial-lateral tilt angle, the medial-lateral rotation angle and the dorsal displacement distance of the talus head of the talus prosthesis in the range of motion of the ankle joint dorsiflexion 5°, dorsiflexion 10°, neutral position, plantar flexion 5°, plantar flexion 10°, plantar flexion 20° and plantar flexion 30° is tested.

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