Biomechanical analysis system for stability and stress distribution of talus prosthesis

Abstract

The invention relates to the technical field of prosthesis mechanical analysis, in particular to a talus prosthesis stability and stress distribution biomechanical analysis system which is composed of a three-dimensional modeling module, a finite element analysis module, a stability analysis module and a report generation module. The three-dimensional modeling module processes CT data through a reverse engineering technology; constructing a three-dimensional model of the talus prosthesis; the finite element analysis module is used for simulating stress conditions of the talus prosthesis at different positions and cartilage covering layers with different thicknesses; the stability analysis module is used for analyzing stress distribution conditions at different positions after ligaments of the talus prosthesis are sequentially cut off through finite element calculation; and the report generation module is used for generating a stress distribution cloud picture and displaying the stress distribution condition of the surface of the talus prosthesis. According to the method, the model of the talus prosthesis is established, stress distribution is predicted based on finite element analysis, and potential failure points or excessive wear are identified, so that design improvement and material selection are guided.

Description

Technical Field

[0001] The invention relates to the technical field of prosthesis mechanics analysis, and in particular to a biomechanical analysis system for stability and stress distribution of a talar prosthesis. Background Art

[0002] The talus is an important bone in the human foot, connecting the lower limbs and the foot. It has an irregular shape, is covered by five articular surfaces, and involves multiple movable joints. It is the mechanical turning point between the lower limbs and the foot. There are abundant ligaments and tendons passing around the talus to maintain its stability and motor function. Due to the unique anatomical and biomechanical properties of the talus, it is prone to ischemic necrosis under factors such as trauma and autoimmune diseases. In severe cases, it can lead to talus collapse, seriously affecting the patient's standing and walking functions.

[0003] In the prior art, joint fusion is usually used to treat talus collapse necrosis. Although this method can relieve pain, it will sacrifice the function of the ankle joint and limit the patient's range of activities in daily life, especially when walking up and down stairs and running. In addition, joint fusion may also cause compensatory degeneration of surrounding joints, as well as muscle atrophy and osteoporosis caused by long-term fixation.

[0004] In summary, how to solve the problem of limiting the range of joint motion when using joint fusion to treat talar collapse necrosis in the prior art has become a difficult problem that needs to be solved urgently in this field. Therefore, it is necessary to propose a biomechanical analysis system for talar prosthesis stability and stress distribution. Summary of the invention

[0005] To solve the above problems, the present invention provides a biomechanical analysis system for talar prosthesis stability and stress distribution. By establishing a model of the talar prosthesis, the stress distribution is predicted based on finite element analysis and potential failure points or excessive wear are identified, thereby guiding design improvements and material selection.

[0006] In order to achieve the above-mentioned purpose, the technical scheme of the present invention is as follows: a biomechanical analysis system for talar prosthesis stability and stress distribution, which consists of a three-dimensional modeling module, a finite element analysis module, a stability analysis module and a report generation module.

[0007] 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 talar prosthesis; and design a geometric model of the talar prosthesis based on the geometric shapes of the talus, tibia and calcaneus.

[0008] The finite element analysis module is used to import the three-dimensional model into the 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 talar prosthesis at different positions and with different thicknesses of cartilage covering layers.

[0009] The stability analysis module is used to analyze the stress distribution of the talar prosthesis at different positions after cutting different ligament combinations in sequence through finite element calculation; the overall stability of the talar prosthesis is evaluated by combining the stress distribution results and the geometric shape of the talar prosthesis.

[0010] The report generation module is used to generate stress distribution cloud maps to display the stress distribution on the surface of the talar prosthesis; and to generate data analysis reports based on the results of finite element analysis calculations.

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

[0012] S1, data acquisition and processing: The original DICOM format data of the CT scan was imported into the Mimics software, and the initial model was obtained by identifying different parts of the bones and cartilage by setting different grayscale values; the talus, tibia and calcaneus were exported as model data files in STL format.

[0013] S2, surface treatment and optimization: Use the "Relax", "Remove Spikes" and "Remove Filtering" commands to fine-tune the initial model surface, retaining the model's features while processing the surface; import the STL file into Geomagic software to 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 interference checking to identify and record overlapping parts.

[0015] S4, create a three-dimensional model: use SolidWorks to create a ligament model, the ligament model includes the lateral, medial, talonavicular and tarsal sinus ligaments; add the skin to the completed ankle joint model, and save the model in X_T format.

[0016] Furthermore, in the finite element analysis module, the material property is titanium alloy, and its elastic modulus is set to 105000MPa and its density is set to 4.5g / cm 3 , the Poisson's ratio is 0.37.

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

[0018] Furthermore, in the finite element analysis module, different positions of the talar prosthesis included neutral, dorsiflexion 5°, dorsiflexion 10°, plantar flexion 5°, plantar flexion 10°, plantar flexion 20°, and plantar flexion 30°.

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

[0020] Furthermore, in the stability analysis module, after cutting off the lateral-medial-tarsal sinus-talonavicular ligament in sequence, the talar prosthesis was tested in the range of motion of 5° dorsiflexion, 10° dorsiflexion, neutral position, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion to test the changes in the anterior-posterior displacement of the talus, the internal and external valgus angle, the internal and external rotation angle and the dorsal displacement distance of the talar head.

[0021] Furthermore, in the stability analysis module, after cutting off the medial-tarsal sinus-talonavicular-lateral ligament in sequence, the talar prosthesis was tested in the range of motion of 5° dorsiflexion, 10° dorsiflexion, neutral position, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion to test the changes in the anterior-posterior displacement of the talus, the internal and external valgus angle, the internal and external rotation angle and the dorsal displacement distance of the talar head.

[0022] Furthermore, in the stability analysis module, after cutting off the tarsal sinus-talonavicular-lateral-medial ligament in sequence, the talar prosthesis was tested in the range of motion of 5° dorsiflexion, 10° dorsiflexion, neutral position, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion to test the changes in the anterior-posterior displacement of the talus, the internal and external valgus angle, the internal and external rotation angle and the dorsal displacement distance of the talar head.

[0023] Furthermore, in the stability analysis module, after the talonavicular-lateral-medial-tarsal sinus ligament was cut off in sequence, the talar prosthesis was tested in the range of motion of 5° dorsiflexion, 10° dorsiflexion, neutral position, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion to test the changes in the anterior-posterior displacement of the talus, the internal and external valgus angle, the internal and external rotation angle and the dorsal displacement distance of the talar head.

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

[0025] 1. The present invention applies finite element analysis to post-placement testing of talar prostheses. The process starts with creating a detailed and accurate model of the ankle joint, including the implanted prosthesis and the surrounding bone and tissue structures. By using high-resolution imaging data and advanced modeling software, an accurate representation of anatomical features can be generated. Insights gained from finite element analysis simulations can provide information for clinical practice. Surgeons can use this information to make more informed decisions about prosthesis selection, placement techniques, and postoperative care. By understanding the biomechanical interactions within the ankle joint, surgical outcomes can be optimized and patient recovery can be promoted.

[0026] In summary, the use of FEA in the evaluation of talar prostheses is a critical step in improving the design and performance of these implants. By providing a detailed understanding of the mechanical behavior of the prosthesis and its interaction with surrounding structures, FEA helps ensure that patients receive a durable, functional, and reliable solution for severe talar injuries. The structural integrity of the prosthesis and its interaction with the surrounding bone and tissue can be evaluated in detail. Second, FEA helps predict stress distribution and identify potential failure points or excessive wear, thereby guiding design improvements and material selection. It provides insights into the performance of the prosthesis under different loading conditions, which is critical to ensuring its durability and effectiveness in restoring function.

[0027] 2. The present invention processes the talar prosthesis by utilizing a variety of finite element analysis software, and identifies and constructs high-precision initial models of the talus, tibia, and calcaneus based on the original data in DICOM format of the CT scan. And through optimization processing, the initial model surface can be fine-tuned to remove unnecessary details and noise while retaining the key features of the model; it helps to simplify the complexity of the model, improve computational efficiency, and maintain the biological authenticity of the model. Through assembly and error checking performed by SolidWorks, overlapping parts are identified and recorded after assembly to prevent biomechanical analysis errors caused by improper contact between parts. Adding soft tissues such as ligaments and skin to the three-dimensional model can more realistically simulate the actual state of the human ankle joint; it helps to more comprehensively evaluate the stability and stress distribution of the talar prosthesis and improve the accuracy and reliability of the analysis.

[0028] 3. The present invention evaluates the influence of the cartilage covering layer on the stability and stress distribution of the prosthesis by analyzing the stress distribution of the talar prosthesis under the cartilage covering layer of different thicknesses of +0.5mm, +1.0mm, +1.5mm and +2.0mm; this helps to optimize the design of the cartilage covering layer and improve the biocompatibility and service life of the prosthesis. By analyzing the stress distribution of the talar prosthesis in different positions such as neutral position, dorsiflexion 5°, dorsiflexion 10°, plantar flexion 5°, plantar flexion 10°, plantar flexion 20° and plantar flexion 30°, it is possible to have a more comprehensive understanding of the mechanical properties and stability of the prosthesis; the performance of the prosthesis in various common physiological activities is comprehensively evaluated, which is of great significance for understanding the stability and stress distribution of the prosthesis in different movement modes.

[0029] 4. The present invention can comprehensively evaluate the stability of the prosthesis by cutting different combinations of lateral, medial, tarsal sinus, and talonavicular ligaments and testing the displacement and angle changes of the talar prosthesis under different ranges of motion of the ankle joint. Comparing the effects of different ligament cutting sequences on the stability of the talar prosthesis can provide a direction for improving the design of the prosthesis.

[0030] Additional aspects and advantages of the present invention will be given in part in the following description and in part will be obvious from the following description, or will be learned through practice of the present invention. BRIEF DESCRIPTION OF THE DRAWINGS

[0031] Figure 1 4 is a structural block diagram of a biomechanical analysis system for talar prosthesis stability and stress distribution in an embodiment of the present invention.

[0032] Figure 2 The figure is a flowchart of the reverse engineering technology processing in the embodiment of the present invention. DETAILED DESCRIPTION

[0033] The following is further described in detail through specific implementation methods:

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

[0035] 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 talar prosthesis; and design a geometric model of the talar prosthesis based on the geometric shapes of the talus, tibia and calcaneus.

[0036] Among them, the reverse engineering technology processing includes the following steps:

[0037] S1, data acquisition and processing: The original DICOM format data of the CT scan was imported into the Mimics software, and the initial model was obtained by identifying different parts of the bones and cartilage by setting different grayscale values; the talus, tibia and calcaneus were exported as model data files in STL format.

[0038] S2, surface treatment and optimization: Use the "Relax", "Remove Spikes" and "Remove Filtering" commands to fine-tune the initial model surface, retaining the model's features while processing the surface; import the STL file into Geomagic software to 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 interference checking to identify and record overlapping parts.

[0040] S4, create a three-dimensional model: use SolidWorks to create a ligament model, the ligament model includes the lateral, medial, talonavicular and tarsal sinus ligaments; add the skin to the completed ankle joint model, and save the model in X_T format.

[0041] The finite element analysis module is used to import the 3D model into the finite element analysis software for meshing, and to apply the material properties to titanium alloy, setting its elastic modulus to 105000MPa and density to 4.5g / cm 3 , Poisson's ratio is 0.37; according to the principles of human physiology and biomechanics, 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; the force conditions of the talar prosthesis in different positions of neutral, dorsiflexion 5°, dorsiflexion 10°, plantar flexion 5°, plantar flexion 10°, plantar flexion 20° and plantar flexion 30° and the stress conditions of the cartilage covering layers of 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 of the talar prosthesis at different positions after cutting different ligament combinations in sequence through finite element calculation; the overall stability of the talar prosthesis is evaluated by combining the stress distribution results and the geometric shape of the talar prosthesis.

[0043] The different ligament combinations were cutting off the lateral-medial-tarsal sinus-talonavicular ligament, medial-tarsal sinus-talonavicular-lateral ligament, tarsal sinus-talonavicular-lateral-medial ligament and talonavicular-lateral-medial-tarsal sinus ligament in sequence, and the talar prosthesis was tested in the range of motion of 5° dorsiflexion, 10° dorsiflexion, neutral position, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion of the ankle joint to test the changes in the anterior-posterior displacement of the talus, the internal and external valgus angle, the internal and external rotation angle and the dorsal displacement distance of the talar head.

[0044] The report generation module is used to generate stress distribution cloud maps to display the stress distribution on the surface of the talar prosthesis; and to generate data analysis reports based on the results of finite element analysis calculations.

[0045] Experiment and demonstration

[0046] 1. Methods

[0047] Data Source

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

[0049] 2. Basic steps to build ankle joint model:

[0050] S1, data acquisition and processing: The original DICOM format data of the CT scan was imported into the Mimics software, and the initial model was obtained by identifying different parts of the bones and cartilage by setting different grayscale values; the talus, tibia and calcaneus were exported as model data files in STL format.

[0051] S2, surface treatment and optimization: Use the "Relax", "Remove Spikes" and "Remove Filtering" commands to fine-tune the initial model surface, retaining the model's features while processing the surface; import the STL file into Geomagic software to 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 interference checking to identify and record overlapping parts.

[0053] S4, create a three-dimensional model: use SolidWorks to create a ligament model, the ligament model includes the lateral, medial, talonavicular and tarsal sinus ligaments; add skin to the completed ankle joint model, save the model in X_T format, and import it into Ansys Workbench for analysis settings.

[0054] After the model was created, the mechanical properties of the bone, prosthesis, skin, ligaments, and cartilage were incorporated into the model. Different ankle joint positions were simulated under different ligament resection conditions, including neutral position, dorsiflexion 5°, dorsiflexion 10°, plantar flexion 5°, plantar flexion 10°, plantar flexion 20°, and plantar flexion 30°. The surrounding ligaments were resected in sequence, including the lateral, medial, sinus tarsi, and talonavicular ligaments. The cartilage covering layers of different thicknesses of +0.5mm, +1.0mm, +1.5mm, and +2.0mm were 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 volume according to the model characteristics. Based on experience, fine-tune the mesh to improve the quality and data accuracy. Apply the material property to titanium alloy. In the static structure module, right-click to insert fixed constraints on the upper ends of the tibia and fibula, and apply a force of 500N to the foot. Insert the results of the maximum equivalent stress and deformation. Select the corresponding model in the parameter settings to extract the required results.

[0057] 4. Discussion of results

[0058] 4.1 talar prosthesis + 0mm

[0059] The overall displacement of the talar component showed differences in different positions (Table 1). In the neutral position, it ranged from 1.2474 mm to 1.7748 mm. The displacement increased during dorsiflexion, reaching 2.0935 mm at 5° and 2.125 mm at 10°. During plantar flexion, the displacement decreased with increasing angle; it was 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 the displacement is most significant during dorsiflexion and is minimal during high plantar flexion.

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

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

[0062] The Z-axis displacement was positive in all positions, indicating upward movement. In the neutral position, it was 0.54856 mm, increasing to 0.6273 mm (5°) and 0.64081 mm (10°) during dorsiflexion. In plantar flexion, the displacement was 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 plantar flexion angle.

[0063] The contact area in the neutral position is the largest, at 480.97 mm2 (Table 1). It is slightly lower when the back is extended, at 480.11 mm 2 (5°) and 479.86mm 2 (10°). In plantar flexion, the contact area remains around 480.94mm 2 (5°) and 480.93mm 2 (10°), but significantly reduced to 474.04mm2 (20°) and 445.35mm at 30° 2These changes indicate that contact is stable at lower flexion angles and more variability at higher angles.

[0064] There were significant differences in contact stress at different positions (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, with a maximum of 2.7272 MPa at 5° and 3.0024 MPa at 10°. In plantar flexion, the maximum stress was 2.0474 MPa and 2.0242 MPa at 5° and 10°, respectively, but increased with increasing angles, reaching 6.9746 MPa at 20° and 6.2178 MPa at 30°. The minimum stress was 0.18389 MPa. These results indicate the effect of flexion angle on stress distribution.

[0065] 4.2 talar prosthesis + 0.5mm

[0066] The total displacement of the talar component with an offset of 0.5 mm was significantly different between different positions (Table 2). In the neutral position, the displacement was 2.2572 mm. During 5° and 10° dorsiflexion, 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 significant during dorsiflexion and 20° plantar flexion, while the displacement was minimal during high plantar flexion.

[0067] The X-axis displacement varied significantly with position (Table 2). In the neutral position, the displacement was -0.05795 mm. During 5° dorsiflexion, this value increased to -0.27465 mm and decreased to -0.14123 mm at 10°. During plantar flexion, the displacement varied 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 showed consistent negative values, indicating downward movement (Table 2). In the neutral position, the displacement was -0.7499 mm. During dorsiflexion at 5° and 10°, it increased to -1.4337 mm and -1.3591 mm, respectively. In plantar flexion, the displacement was -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 flexion angle on the prosthesis.

[0069] The Z-axis displacement showed a positive trend in all positions, indicating an upward movement (Table 2). In the neutral position, the displacement was 0.33928 mm. In 5° and 10° dorsiflexion, the displacements were 0.62143 mm and 0.60524 mm, respectively. In plantar flexion, the displacements were 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 flexion angle and is affected by the direction of movement.

[0070] The contact area in the neutral position is the largest, at 481.64mm 2 (Table 2). During dorsiflexion, it decreased slightly to 480.96 mm at 5° and 10°, respectively. 2 and 472.14mm 2 In plantar flexion, the contact area remained relatively stable at 468.47 mm at 5° and 10°, respectively. 2 and 469.44mm 2 , but when it is reduced to 20°, it is 474.04mm 2 ,445.24mm at 30° 2 These changes indicate that contact is stable at lower flexion angles and more variability at higher angles.

[0071] The contact stress experienced by the talar prosthesis in different positions was different (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, with a maximum of 2.6961 MPa at 5° and 2.7708 MPa at 10°, and the corresponding minimum values ​​were 0.32898 MPa and 0.24591 MPa. The maximum stress in plantar flexion was 2.0466 MPa at 5°, 1.459 MPa at 10°, 2.434 MPa at 20°, and 1.1884 MPa at 30°, and the minimum value was 0.17031 MPa. These results show that different flexion angles affect stress distribution, and higher angles lead to increased stress.

[0072] 4.3 talar prosthesis + 1mm

[0073] The overall displacement of the 1 mm offset talar component showed substantial differences between positions (Table 3). In the neutral position, the displacement was 1.7882 mm. During 5° and 10° dorsiflexion, 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 suggests that the displacement is most significant during dorsiflexion and is minimal during higher degrees of plantar flexion.

[0074] The X-axis displacement varied significantly with position (Table 3). In the neutral position, the displacement was -0.12261 mm. During 5° dorsiflexion, this value increased to -0.26078 mm and decreased to -0.15449 mm at 10°. During plantar flexion, the displacement ranged 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 showed consistent negative values, indicating downward movement (Table 3). In the neutral position, the displacement was -1.2384 mm. During dorsiflexion at 5° and 10°, it increased to -1.4001 mm and -1.4614 mm, respectively. In plantar flexion, the displacement was -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 flexion angle on the prosthesis.

[0076] The Z-axis displacement showed a positive trend in all positions, indicating an upward movement (Table 3). In the neutral position, the displacement was 0.56644 mm. In 5° and 10° dorsiflexion, it was 0.59316 mm and 0.6469 mm, respectively. In plantar flexion, the displacement was 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 flexion angle and is affected by the direction of movement.

[0077] The contact area in the neutral position is the largest, at 482.82mm 2 (Table 3). During dorsiflexion, it decreased slightly to 480.96 mm at 5° and 10°, respectively. 2 and 468.42mm 2 In plantar flexion, the contact area remained relatively stable at 460.74 mm at 5° and 10°, respectively. 2 and 465.23mm 2 , but significantly reduced to 474.04mm at 20° 2 and 30°: 445.47mm 2 These variations indicate stable contacts at lower flexion angles and greater variability at higher angles. The talar component experiences different contact stresses at different positions (Table 3).

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

[0079] 4.4 talar prosthesis + 1.5mm

[0080] The total displacement of the 1.5 mm offset talar component was significantly different in different positions (Table 4). In the neutral position, the displacement was 1.9302 mm. During 5° and 10° dorsiflexion, 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 is most pronounced during dorsiflexion and is minimal during higher degrees of plantar flexion.

[0081] The X-axis displacement varied significantly with position (Table 4). In the neutral position, the displacement was -0.20596 mm. In dorsiflexion, it was -0.25359 mm at 5° and -0.19579 mm at 10°. In plantar flexion, the displacement ranged 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 showed consistent negative values, indicating downward movement (Table 4). In the neutral position, the displacement was -1.326 mm. During dorsiflexion at 5° and 10°, it increased to -1.4455 mm and -1.4409 mm, respectively. In plantar flexion, the displacement was -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 flexion angle on the prosthesis.

[0083] The Z-axis displacement showed a positive trend in all positions, indicating an upward movement (Table 4). In the neutral position, the displacement was 0.58021 mm. In 5° and 10° dorsiflexion, it was 0.6327 mm and 0.62447 mm, respectively. In plantar flexion, the displacement was 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 flexion angle and is affected by the direction of movement.

[0084] The contact area in the neutral position is the largest, at 481.64mm 2 (Table 4). During dorsiflexion, it decreased slightly to 480.97 mm at 5° and 10°, respectively. 2 and 470.55mm 2 In plantar flexion, the contact area remained relatively stable at 5° and 10°, at 468.43 mm and 10.3 mm, respectively. 2 and 469.46mm 2 , but it drops to 474.04mm at 20° and 30° respectively 2 and 456.99mm 2 These changes indicate that the contact is stable at lower flexion angles and more variable at higher angles. The talar component experiences different contact stresses at different positions (Table 4).

[0085] In the neutral position, the maximum contact stress is 3.0293MPa and the minimum is 0.3016MPa. During the dorsiflexion process, the stress increases, with the maximum value of 2.7719MPa at 5° and 3.2242MPa at 10°, and the corresponding minimum values ​​are 0.29949MPa and 0.27412MPa, respectively. In plantar flexion, the maximum stress is 2.2429MPa at 5°, 1.7929MPa at 10°, and 1.3392MPa at 10°. At 30°, it is 1.0711MPa, and the minimum value is 0.17669MPa. These results show the influence of different flexion angles on stress distribution, and the larger the angle, the greater the stress.

[0086] 4.5 talar prosthesis + 2mm

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

[0088] The X-axis displacement varied significantly with position (Table 5). In the neutral position, the displacement was -0.16778 mm. During dorsiflexion, it was -0.20529 mm at 5° and -0.15239 mm at 10°. In plantar flexion, the displacement was -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 plantar flexion angle indicates lateral movement of the prosthesis.

[0089] The Y-axis displacement showed consistent negative values, indicating downward movement (Table 5). In the neutral position, the displacement was -1.3287 mm. During dorsiflexion at 5° and 10°, it increased to -1.5315 mm and -1.4535 mm, respectively. In plantar flexion, the displacement was -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 showed a positive trend in all positions, indicating an upward movement (Table 5). In the neutral position, the displacement was 0.58314 mm. In 5° and 10° dorsiflexion, the displacements were 0.67009 mm and 0.63508 mm, respectively. In plantar flexion, the displacements were 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 flexion angle and is affected by the direction of movement.

[0091] The contact area in the neutral position is the largest, at 481.63mm 2 (Table 5). During dorsiflexion, it decreased slightly to 480.94 mm at 5° and 10°, respectively. 2 and 468.56mm 2 In plantar flexion, the contact area remained relatively stable at 468.44 mm at 5° and 10°, respectively. 2 and 469.48mm 2 , but significantly reduced to 474.84mm at 20° and 30° 2 and 458.05mm 2 These changes indicate that the contact is stable at lower flexion angles and more variable at higher angles. The talar component experiences different contact stresses at different positions (Table 5).

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

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

[0094]

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

[0096]

[0097]

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

[0099]

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

[0101]

[0102]

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

[0104]

[0105] Obviously, the above embodiments are merely examples for the purpose of clear explanation, and are not intended to limit the implementation methods. For those skilled in the art, other different forms of changes or modifications can be made based on the above description. It is not necessary and impossible to list all the implementation methods here. The obvious changes or modifications derived therefrom are still within the scope of protection of the invention.

Claims

1. A biomechanical analysis system for stability and stress distribution of talar prosthesis, 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, and process the CT data through reverse engineering technology; construct a three-dimensional model of the talar prosthesis; and design a geometric model of the talar prosthesis according to the geometric shapes of the talus, tibia and calcaneus; The finite element analysis module is used to import the 3D model into the finite element analysis software for meshing and applying material properties; Boundary conditions and loads were set according to human physiology and biomechanics to simulate the stress of the talar prosthesis at different positions and with different thicknesses of cartilage covering layers; The stability analysis module is used to analyze the stress distribution of the talar prosthesis at different positions after cutting off different ligament combinations in sequence through finite element calculation; the overall stability of the talar prosthesis is evaluated by combining the stress distribution results and the geometric shape of the talar prosthesis; The report generation module is used to generate stress distribution cloud maps to display the stress distribution on the surface of the talar prosthesis; and to generate data analysis reports based on the results of finite element analysis calculations.

2. The biomechanical analysis system for talar prosthesis stability and stress distribution according to claim 1, characterized in that: In the 3D modeling module, the reverse engineering process includes the following steps: S1, data acquisition and processing: import the original DICOM format data of CT scan into Mimics software, and identify different parts of bones and cartilage by setting different gray values ​​to obtain the initial model; export the talus, tibia and calcaneus into STL format model data files; S2, surface treatment and optimization: Use the "relax", "remove spikes" and "remove filters" commands to fine-tune the initial model surface, preserving the model's features while processing the surface; 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 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 interference checking to identify and record overlapping parts; S4, create a three-dimensional model: use SolidWorks to create a ligament model, the ligament model includes the lateral, medial, talonavicular and tarsal sinus ligaments; add the skin to the completed ankle joint model, and save the model in X_T format.

3. The biomechanical analysis system for talar prosthesis stability and stress distribution according to claim 2, characterized in that: In the finite element analysis module, the material property is titanium alloy, and its elastic modulus is set to 105000MPa and its density is set to 4.5g / cm 3 , the Poisson's ratio is 0.

37.

4. The biomechanical analysis system for talar prosthesis stability and stress distribution according to claim 3, characterized in that: In the finite element analysis module, the boundary conditions were set to insert fixed constraints at the upper ends of the tibia and fibula, and the load was set to apply a force of 500N to the foot.

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

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

7. The biomechanical analysis system for talar prosthesis stability and stress distribution according to claim 6, characterized in that: In the stability analysis module, the lateral-medial-tarsal sinus-talonavicular ligament was cut off in sequence, and the talar prosthesis was tested in the range of motion of 5° dorsiflexion, 10° dorsiflexion, neutral position, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion to test the changes in the anterior-posterior displacement of the talus, the internal and external valgus angle, the internal and external rotation angle, and the dorsal displacement distance of the talar head.

8. The biomechanical analysis system for talar prosthesis stability and stress distribution according to claim 6, characterized in that: In the stability analysis module, the medial-tarsal sinus-talonavicular-lateral ligament was cut off in sequence, and the talar prosthesis was tested in the range of motion of 5° dorsiflexion, 10° dorsiflexion, neutral position, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion to test the changes in the anterior-posterior displacement of the talus, the internal and external valgus angle, the internal and external rotation angle, and the dorsal displacement distance of the talar head.

9. The biomechanical analysis system for talar prosthesis stability and stress distribution according to claim 6, characterized in that: In the stability analysis module, the tarsal sinus-talonavicular-lateral-medial ligament was cut off in sequence, and the talar prosthesis was tested in the range of motion of 5° dorsiflexion, 10° dorsiflexion, neutral position, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion to test the changes in the anterior-posterior displacement of the talus, the internal and external valgus angle, the internal and external rotation angle, and the dorsal displacement distance of the talar head.

10. The biomechanical analysis system for talar prosthesis stability and stress distribution according to claim 6, characterized in that: In the stability analysis module, the talonavicular-lateral-medial-tarsal sinus ligament was cut off in sequence, and the talar prosthesis was tested in the range of motion of 5° dorsiflexion, 10° dorsiflexion, neutral position, 5° plantar flexion, 10° plantar flexion, 20° plantar flexion and 30° plantar flexion to test the changes in the anterior-posterior displacement of the talus, the internal and external valgus angle, the internal and external rotation angle, and the dorsal displacement distance of the talar head.

Images