Issues of Knee Joint Endoprosthetics in Megadefects Cases: Opportunities and Challenges

KEYWORDS
knee prosthesis, revision prosthetics, Solidworks, Simsolid, finite element modeling

ABSTRACT

Total knee arthroplasty (TKA) is an effective surgical method aimed at relieving pain and restoring joint function in patients with degenerative diseases and post-traumatic conditions. However, over time, many patients require revision procedures. Revision TKA is technically more demanding than primary arthroplasty and is associated with less predictable outcomes and a higher risk of complications, mainly due to extensive bone and soft tissue defects. Management of massive bone defects remains a major challenge and requires optimization of implant design and fixation strategies.

Objective. – The aim of this study was to perform a computational analysis of the biomechanical performance of individualized hinge-type knee prostheses compared with standardized implants in the presence of large bone defects.

Design. – Finite element modeling was used to simulate revision knee arthroplasty under conditions of extensive bone loss. The inclination angle of intramedullary extensions was set at 5° according to the anatomical axis of the bone. Three lengths of individualized extensions (120 mm, 130 mm, and 150 mm) were evaluated. Stress distribution, displacement, deformation, and factor of safety (FoS) were analyzed for individualized and standard prosthetic configurations.

Results. – The 130 mm individualized extension demonstrated the most favorable biomechanical performance, including minimal displacement, reduced stress concentration, and sufficient safety margin. Individually designed augments showed no adverse mechanical effects. Standardized prosthetic components produced excessive stress in residual bone exceeding 3.5 MPa, significant deformation, and insufficient structural safety (FoS < 1).

Conclusions. – The use of computer navigation technologies and individually designed components allows for increased precision in endoprosthesis placement, improved biomechanical compatibility, and a reduced risk of damage to residual bone tissue in the postoperative period.

ARTICLE HISTORY
Received November 14, 2025
Accepted December 15, 2025

Introduction

Bone mega-defects of the knee joint remain one of the most difficult problems of modern orthopedics and reconstructive arthroplasty. Most often, they are formed due to aseptic loosening of the endoprosthesis components, periprosthetic infection, repeated revision interventions, periprosthetic fractures,post-traumatic injuries and significant loss of metaphyseal bone mass. 1,2.

In clinical practice, such defects are especially characteristic of revision knee arthroplasty, where insufficient bone support directly affects the possibility of restoring the limb axis, implant stability and long-term survival of the reconstruction. Modern reviews emphasize that bone deficiency is one of the key factors that complicate revision knee arthroplasty and worsen functional results compared to primary arthroplasty. 3,4,5 Often, megadefects of the femoral and tibial condyles can occur in young patients with complications after reconstruction of the anterior cruciate ligament, including among military personnel. These defects occur after reactive arthritis, synovitis of the knee joint after reconstructive and restorative surgical interventions, as well as in patients with expansion of bone tunnels as a result of bioresorbable implants lysis or infectious process.

The urgency of this problem is increasing in parallel with increase in the number of primary and revision arthroplasties. According to current epidemiological forecasts, the burden on the health care system associated with revision knee arthroplasty will continue to increase: the expected increase in the number of revision TKAs in the United States is about 149% by 2040 and 520% by 2060.6 This means that the number of patients with severe metaphyseal and metaphyseal defects will also increase, and the problem of choosing the optimal reconstructive tactics will become even more important. It should be noted that infectious complications remain one of the leading causes of revisions, and with repeated interventions, the frequency of periprosthetic infection is significantly higher than with primary arthroplasty, which further deepens bone loss and complicates reconstruction. 7,8

In recent years, researchers focused attention on improving methods for replacing large bone defects in revision knee arthroplasty. 9 The most common modern solutions are metaphyseal cones and sleeves, which allow for reliable metaphyseal fixation in conditions of significant bone loss. Systematic reviews, published in recent years, demonstrate that both technologies show good clinical and radiographic results, satisfactory implant survival, and can be considered effective tools for severe defects, although there is still no universal “gold standard”. 10,11 At the same time, cones and sleeves do not enable solving all problems, especially in patients with AORI III type defects, severe anatomical abnormalities, lack of bone support, or the consequences of infectious and traumatic lesions. 12,13,14

In the most challenging cases, where standard modular solutions do not provide full bone contact or do not allow for the restoration of segmental biomechanics, personalized 3D-printed implants are becoming increasingly important. According to clinical series, the use of personalized highly porous 3D-printed metaphyseal cones in massive proximal tibial defects yielded encouraging clinical and radiographic results with no evidence of aseptic loosening or reinfection in short- and medium-term follow-up. 15,16,17,18. This approach is particularly promising in patients with non-standard defect geometry, extensive segmental bone loss, and complex post-traumatic or post-infectious changes.

However, in extreme mega-defects, where bone loss exceeds the possibilities of metaphyseal reconstruction, salvage treatment options, such as resection arthroplasty, distal femoral replacement, or megaprosthetics, must be considered19. Such interventions can ensure limb preservation and support, but are accompanied by high complexity, a significant risk of reoperation, and significant requirements for preoperative planning3. Thus, the problem of mega-defects of the knee joint today goes far beyond simple technical bone reconstruction and includes issues of biomechanics, infection control, individualization of implants, and the selection of the optimal fixation strategy.

Thus, mega-defects of the knee joint represent one of the most urgent and not yet fully resolved problems of revision orthopedics. The increase in the number of revision interventions, the high frequency of infectious and mechanical complications, the limitations of standard reconstructive solutions and the emergence of new additive technologies determine the need for further study of effective approaches to the replacement of large bone defects, especially in patients with defects of a high degree of complexity.

In parallel with the development of reconstructive strategies, increasing attention was paid to biomechanical optimization of revision knee prostheses using digital simulation technologies. Finite element analysis (FEA) has become one of the key tools for predicting implant behavior in the presence of severe bone loss, as it enables detailed assessment of stress distribution, micromotion, and deformation patterns under physiological loading conditions. Recent biomechanical studies demonstrated that computer-based modeling significantly improves understanding of load transfer mechanisms in revision TKA and allows identification of critical stress concentration zones that may predispose to implant loosening or bone resorption20,21. Such modeling approaches are particularly valuable in cases of AORI type III defects, where the remaining bone stock is limited and conventional fixation techniques may not provide sufficient mechanical stability.

Another important direction in modern reconstructive orthopedics is the use of patient-specific implants and individualized fixation strategies. With the advancement of additive manufacturing technologies, custom-designed prosthetic components have become increasingly accessible and clinically feasible. Recent clinical and biomechanical investigations indicate that individualized prosthetic systems allow improved conformity between implant geometry and residual bone anatomy, thereby enhancing stress distribution and reducing peak mechanical loads at the bone–implant interface. 22,15 This is particularly relevant for patients with irregular defect morphology, where standard modular systems may lead to suboptimal contact conditions and increased risk of implant migration or instability.

In addition to implant geometry, the length and configuration of intramedullary stems remain critical parameters influencing prosthetic stability in revision knee arthroplasty. Several contemporary biomechanical investigations have demonstrated that optimization of stem length significantly affects load transfer patterns and micromotion at the bone–implant interface. Excessively short stems may lead to insufficient fixation, whereas overly long stems can generate stress shielding or create unfavorable mechanical leverage, increasing the risk of bone damage. 23,17 Therefore, the selection of optimal stem dimensions represents a key biomechanical challenge, particularly in cases of massive metaphyseal and diaphyseal bone defects.

Despite the growing body of literature addressing revision knee arthroplasty techniques, there remains a lack of consensus regarding the optimal configuration of prosthetic components in patients with large-scale bone defects. Most existing studies focus on clinical outcomes or isolated mechanical parameters, whereas comprehensive biomechanical analyses that integrate defect geometry, implant design, and loading conditions remain relatively limited. Furthermore, comparative investigations evaluating individualized and standardized prosthetic systems under identical biomechanical conditions are still insufficient, highlighting the need for further research in this field.

With this in mind, the study aims to determine the conditions of knee joint endoprosthesis in patients with bone defects by simulating an anthropometric model, and to assess the expected impact of knee joint endoprosthesis components on residual bone tissue in patients with megadefects.

Materials and Methods

This study represents a computational biomechanical simulation investigation aimed at evaluating the stress–strain behavior of knee joint prosthetic systems in the presence of large bone defects corresponding to AORI type III classification. The modeling approach was selected due to its ability to simulate biomechanical interactions between bone and implant components under controlled loading conditions and to predict potential zones of mechanical failure prior to clinical implementation.

Computer modeling was applied using 3Matic software products (Materialise NV, Leuven, Belgium), Solidworks 2023 (Dassault Systèmes SOLIDWORKS Corp, France), Simsolid 2023 (Altair Engineering Inc, USA).

The anatomical models of the femur and tibia were reconstructed using segmented tomographic data. The primary dataset consisted of computed tomography (CT) images of the knee joint with sufficient spatial resolution to ensure accurate differentiation of cortical and cancellous bone layers. Segmentation of bone structures was performed using the specialized software package 3Matic (Materialise NV, Leuven, Belgium). After segmentation, the resulting three-dimensional models were converted into stereolithography (STL) format and subjected to geometric processing, including smoothing of irregular surfaces and elimination of segmentation artifacts.

Defect modeling corresponding to AORI type III bone defects was performed by selectively removing metaphyseal bone volumes from the femoral and tibial regions. This approach enabled simulation of clinically realistic massive bone deficiencies encountered during revision knee arthroplasty.

Patient-specific digital modeling based on CT-derived geometry has been shown to improve the biomechanical relevance of simulations and enhance prediction accuracy of implant fixation stability.24

The customized prosthesis included a hinge-type knee joint mechanism and a conical intramedullary extender with longitudinal anti-rotation grooves and a porous surface structure designed to facilitate osseointegration. The porous layer thickness ranged between 50 and 100 μm, corresponding to values considered optimal for bone ingrowth and mechanical stabilization. The standardized prosthesis included modular titanium augments with a height of 20 mm, commonly used in revision knee arthroplasty. Individual augments were designed according to the geometry of the simulated bone defects to ensure maximal contact between implant components and residual bone tissue. The geometric accuracy of implant reconstruction was maintained within a tolerance of ±0.02 mm, ensuring consistency between digital models and real prosthetic components. Previous biomechanical investigations have confirmed that customized implant geometries can significantly reduce localized stress concentrations and improve load transfer efficiency compared with standard modular solutions. 21,22.

Finite element modeling

Finite element analysis (FEA) was performed using SimSolid 2023 (Altair Engineering Inc., USA). This software enables accurate simulation of mechanical interactions between multiple components without requiring extensive mesh refinement, which is particularly advantageous for complex implant geometries. Computational models included the following structural components: femoral bone; tibial bone; polyethylene insert; femoral component; tibial component; intramedullary extenders; augmentation elements.

Contact interactions between components were defined as sliding contacts with friction. For the bone–implant interface, friction coefficients ranging from μ = 0.2 to 0.4 were assigned. For the metal–polyethylene interface, friction coefficients ranging from μ = 0.04 to 0.1 were used, reflecting typical operational characteristics of knee prosthetic joints. The osseointegration process was simulated by gradually modifying the contact behavior between the implant and bone from frictional interaction to rigid bonding conditions. Micromotion thresholds exceeding 100–150 μm were considered critical, as such values are associated with fibrous tissue formation rather than bone ingrowth.

Recent finite element studies have emphasized the importance of accurate contact modeling and micromotion analysis in predicting implant longevity and preventing aseptic loosening. 15,17

Loading conditions and boundary parameters

Physiological loading conditions were simulated based on an average body mass of 90 kg, corresponding to dynamic loads ranging from 2 to 3 body weights (BW). The applied force ranged from 1766 N to 2649 N, representing typical peak loading values observed during normal walking cycles.

The loading protocol reproduced key phases of gait, including:

• – heel strike (0–5° flexion)
• – loading response (15–20° flexion)
• – pre-swing phase (30–40° flexion)
• – swing phase (up to 60° flexion)

The total number of loading cycles was set to 30 million, corresponding to long-term functional operation of knee prosthetic systems.

The distal portion of the tibia was rigidly fixed, while vertical loads were applied to the femoral component. Muscle forces were not included in the simulation model to simplify loading conditions and focus on axial load transfer mechanisms.

Modern biomechanical modeling studies confirm that simulation of cyclic loading under physiological conditions enables reliable prediction of fatigue resistance and structural durability of revision prostheses. 23

Outcome assessment parameters

The mechanical performance of prosthetic systems was evaluated using the following key parameters: total displacement of implant components; von Mises equivalent stress; equivalent strain; micromotion at the bone–implant interface; safety factor; reaction and contact forces. These parameters were selected because they reflect the principal biomechanical factors influencing implant survival and structural stability. Particular attention was paid to identifying zones of stress concentration exceeding physiological limits and to evaluating the influence of extender length on the mechanical stability of the bone–implant system.

Recent biomechanical studies demonstrated that optimization of stem length and implant geometry significantly reduces peak stress values and improves long-term fixation stability in revision knee arthroplasty. 25

Results

Paraarticular pain, joint contractures, and aseptic instability of the endoprosthesis components occur in 3% to 12% of surgical interventions. 26 After the first endoprosthesis of the knee joint, bone defects can result from intra- and periarticular fractures of the femoral and tibial condyles. As a result, persistent instability of the prosthesis components is formed, which in a dynamic state has negative progress. When valgus or varus deformities of the knee joint occur, the load on the condyles of the tibia increases significantly, with subsequent axial deformation and subsidence.

Based on the AORI bone defect classification (Fig. 1), 27 grade III defects were generated in Solidworks (Dassault Systèmes SOLIDWORKS Corp, France) in the femur and tibia, the solid-state model of which was obtained from segmented tomographic images, using the 3Matic program (Materialise NV, Leuven, Belgium).

 

Fig. 1. AORI classification of bone defects 28

To process the obtained bone models, the licensed software product Solidworks (Dassault Systèmes SOLIDWORKS Corp, France) was used (Fig. 2.)

 

Fig. 2. Model of femur and tibia with grade III defect

 According to the computer analysis conducted, the volume of the type III defect of the femur is 70.75 cm3, which is equal to 106.125 g. In percentage terms, the type III defect for the femur accounts for 26.53% of its total mass. The type III defect for the tibia is mathematically approximated by a paraboloid of rotation formed by rotating the curve y²=4x² around the abscissa axis (OX), limited by the conditions y=0 and x=3.4.

We apply the limit integral with the limits, a=0; b=3.4 (see (1))

 

Then, the volume of the type III defect is 72.64 cm3, which is 108.9 g., of the total bone mass. In percentage terms, the type III defect for the tibia represents 36.29%.

In accordance with the physical dimensions of the revision prosthesis of one of the leading companies, a 3D model of a standardized prosthesis and an individually designed one was created, the accuracy error is ± 0.02 mm (see Fig. 3.)

 

Fig. 3. 3D model of revision standard and individual knee prostheses

The created model of customized prosthesis has a hinge for the knee joint, which is designed to restore knee function after injury or disease. The individual intramedullary extender has longitudinal grooves/ribs, a modular connection on top, a conical shape, a porous coating for osseointegration of 50-100 microns. The intramedullary rod plays a key role in ensuring the primary and secondary stability of the implant, acting as the main element of load transfer from the prosthesis to the bone tissue. Due to the conical geometry, a self-welding effect (press-fit) is implemented, contributing to tight contact with the inner surface of the bone-medullary canal and a reduction in micromovements in the early postoperative period.

Longitudinal grooves (flutes) perform an anti-rotation function, preventing rotation of the rod relative to the bone axis, which is especially important in the absence of a ligamentous apparatus when using hinge structures. In addition, this geometry contributes to a more uniform distribution of stresses along the rod and a reduction in peak concentrations in the areas of contact “implant-bone”.

The porous coating provides conditions for osseointegration, promoting the germination of bone tissue into the micropores of the implant surface and the formation of a stable biomechanical connection. In combination with the geometry of the rod, this allows achieving a transition from primary mechanical stability to long-term biological fixation.

Thus, the intramedullary extender as part of an individual prosthesis performs not only a fixing, but also a load-distributing function, ensuring effective force transmission, reducing the risk of aseptic loosening and increasing the durability of the entire structure. It allows restoring the range of motion and maintain joint stability. These prostheses have a built-in mechanical loop structure that enables precise control of the knee joint’ flexion and extension angle. For the standard type of revision knee prosthesis, titanium augments with a height of 20 mm are used, and individual augments for the tibia and femur are also developed for a specific condition (Fig. 4).

 

Fig. 4. Standard and custom knee augmentation

The angle of inclination of the extender in the sagittal and coronal planes is 5 degrees, which is relative to the angle of inclination of the bones. Based on the fact that the knee joint with a grade III bone defect is considered, extenders with lengths of 120, 130 and 150 mm were selected for analysis.

The finite element analysis simulation used the Simsolid software product (Altair Engineering Inc, USA).

The initial boundary conditions of the modeling were a body mass of 90 kg, which corresponded to a load at the level of 2–3 body masses (BW), i.e., 1766–2649 N (1.8–2.6 kN), applied to the femur in a vertical downward direction. This range of forces reflects the physiological loading conditions during normal walking and corresponds to peak values in the load acceptance phase.

The model reproduced the main phases of walking with the corresponding angles of knee flexion: heel strike — 0–5°, loading response — 15–20°, in which the main shock absorption occurs, pre-swing — 30–40°, and swing phase — up to 60° of flexion. Thus, the dynamic analysis covered the range of motion from 0° to 60°, which is representative of the full walking cycle without taking into account extreme loads. The number of loading cycles was 30 million,28 which corresponds to long-term operation of the implant and is used as a standard estimate of the fatigue durability of structures.

Contact interactions in the model were specified as Sliding contact with friction. For the bone–implant interface, typical friction coefficients μ = 0.2–0.4 were assumed, reflecting the rough surface and the conditions of primary stabilization. For the metal–polyethylene pair, the friction coefficient was 0.04–0.1, which corresponds to the operating conditions of the hinge assembly of the endoprosthesis and ensures reduced wear.

Bone tissue was modeled as a nonlinear isotropic material using a bilinear stress–strain relationship that accounts for both elastic and plastic behavior after the yield point (Table 1). This approach enables more adequate representation of the real mechanical response of bone under physiological and increased loads, particularly under conditions of local stress concentrations.

The osseointegration of a custom intramedullary extender was modeled by introducing an interaction zone with the rod penetrating (interference) into the bone tissue by 50–100 μm. This range is critically important for the formation of stable bone ingrowth, as it provides a sufficient level of initial mechanical stability and limits micromotion. Osseointegration was modeled by varying the contact conditions between the implant and the bone: from frictional contact (μ = 0.4), which corresponds to initial stability, to a rigid connection (bonded), which simulates complete osseointegration. The evaluation criterion was the magnitude of micromotions in the contact zone. It is known that exceeding this threshold (>100–150 μm) can lead to the formation of fibrous tissue instead of bone, which significantly increases the risk of aseptic loosening of the implant. 29

Table 1. Nonlinearity of the cancellous bone layer

 

Additionally, the model takes into account that the load transfer is predominantly axial in nature with the superposition of bending and torsional components that occur during different phases of walking. This allows not only assessing the overall level of the stress-strain state, but also identifying potentially critical stress concentration zones in the “bone-implant” system. This approach increases the reliability of the prediction of the durability of the structure and the risks of its aseptic instability under conditions of prolonged cyclic loading.

The environment surrounding the prosthetic joint is muscles, the density of which is ρ=1090 kg/m3. Muscle force is neglected. In the lower part, the bone was rigidly fixed. The analysis was performed using Quasi-static, which makes it possible to evaluate the knee joint in dynamics. The system’ behavior was assessed using five key parameters: deformation; micromotion; stress; strain; safety factor; reaction/contact force.

The physical properties of bone tissue correspond to the data presented in Table 2.

Table 2. Physical properties of bone tissue and endoprosthesis elements

When choosing the length of the extension for knee joint prosthetics, it is necessary to consider the individual characteristics of the patient, as well as the features of the pathology and the treatment process, taking into account the size and condition of the patient’s lower limb. It is important to remember that individualized selection of the extension length plays an important role in ensuring optimal comfort and efficiency when using a knee prosthesis. The prosthesis with individually designed augments and with extension lengths of 120, 130 and 150 mm was considered first.

After applying a mass of 90 kg, bone tissue and implant are displaced (Fig. 5).

Fig. 5. Movement (displacement) of bone tissue and components of a custom implant

Comparing the displacement of the prosthetic knee joint, it is noticeable that when using individual revision components with a 120 mm extension, the maximum displacement of the femur is 4.83 mm, when using 130 mm, the displacement is 1.85 mm, and with a 150 mm extension, it is 9.65 mm. For a prosthesis with a 150 mm extension, the displacement is critical, which provokes a possible exit beyond the boundaries of the femur. In all cases, the tibia remains in a stable state, no significant displacements are noted.

The process of modeling and implantation of a custom-made femur augment for knee arthroplasty is depicted in Fig. 6.

 

Fig. 6. Process of modeling and implantation of a custom-made femur augment for knee arthroplasty

After applying a mass of 90 kg to the prosthetic joint, tension arises in the bone tissue (Fig. 7).

Fig. 7. The stress of a custom prosthesis when a mass of 90 kg is applied

The von Mises maximum stress criterion for a knee joint with a 120 mm extender is 2.26 MPa, for a 130 mm extender it is 1.57 MPa, and for a 150 mm extender it is 2.71 MPa. In the first case, the 120 mm extender creates pressure on the femur, thereby causing bone destruction at the interface. This stress and displacement negatively affect the individual augment, causing mechanical bone lysis at the point of its contact with the tibia. An implant with a 130 mm extender creates stress in the femur of 1.57 MPa. The augment and tibia are stable. Use of a revision prosthesis with a 150 mm extender is prone to the formation of a femoral gap at the interface between the extender and the bone. The stress generated when using a 150 mm extension is 2.71 MPa, which leads to changes in the tibia beneath the augment, specifically bone degradation.

Applying 90 kg weight to a static model of a knee replacement results in deformations that cause various types of deviations (see Fig. 8).

Fig. 8. Deformation of the prosthetic knee joint

For a revision knee implant with an extension length of 120 mm, the deformation value is 12.85×10-3. With this type of deformation, its maximum falls on the femur in the extension contact zone, there is a distortion of the shape and destruction of the polyethylene lining and bone tissue at the site of the medial cartilage under the titanium augment. The bone tissue is at the strength limit. For a 130 mm extension, the maximum deformation is 8.96×10-3, there are no noticeable changes in the polyethylene insert, the femur and tibia are stable, without noticeable excessive deformations. In the case of using a prosthesis with an extension length of 150 mm, the deformation that occurs has a value of 15.3977×10-3, the maximum falls on the tibia and femur. According to computer simulation, there is a prerequisite for a fracture of the femur, which can create additional valgus deformity in the tibia.

Simulation computer analysis allows concluding that in revision knee joint prosthetics with grade III bone tissue defects (AORI), the use of extenders with lengths of 120 mm and 150 mm is not advisable. In these cases, significant displacements and overloads were found, provoking bone tissue lysis. Among these studies, a knee implant with an extender with a length of 130 mm should be highlighted. When using this extender, the bone tissue and the augment are unchanged, and the deformation of the polyethylene lining is noticeable, which is within the permissible value. The process of modeling and implantation of a personalized porous titanium augment in endoprosthetics of patients with megadefects is shown in Fig. 9.

 

Fig. 9. Process of modeling and implantation of a personalized porous titanium augment in endoprosthetics of patients with megadefects

Based on the results obtained, the next step is to study a standard revision knee prosthesis and augments with an extender length of 130 mm. The boundary conditions and materials are the same as in the case of the previous study. First, the movement of bone tissue under the action of a mass of 90 kg was analyzed (Fig. 10).

Fig. 10. Movement (displacement) of bone tissue and components of a standard implant

For a standard prosthesis with a 130 mm extender length and standard augments, the maximum displacement is 12.5 mm. Extreme displacement is concentrated at the ends of the extender, with the extenders extending beyond the femur and tibia. This displacement can be explained by the complete absence of ligaments and auxiliary pseudophakic implant connections that stabilize the prosthetic joint. To predict the onset of material damage and bone destruction, a Von Mises stress analysis was performed on the stresses generated by applying a 90 kg mass (Fig. 11).

 

Fig. 11. Tension of a standard prosthesis when applying a mass of 90 kg

Based on this study, the intensity of the external load and the internal resistance of the material were determined. Stress occurs in the lower part of the tibial augment, which is equal to 4.39 MPa, and the lower part of the tibia. Local stresses are noticeable, which provoke a fracture of the bone tissue at the point of contact of the end points of the extender. From the parallel part of the extender fracture, a significant stress was detected, which, according to the infographic, exceeds the permissible norms. The thigh and tibia are subjected to a load along the entire length, ranging from 2.85 MPa to a maximum, which exceeds the strength limit of 3.5 MPa. Due to the violation of the stability of the augments, a significant portion of the stress, 4.39 MPa, falls on the polyethylene overlay. When comparing the stability of the augments, the thigh augment is more or less affected, while the tibia creates a risk of bone tissue destruction.

To determine the change in the shape and structure of a standard prosthetic joint with a 130 mm long extension, under the action of a mass of 90 kg, a computer analysis (Equivalent Strain) was performed (Fig. 12).

Fig. 12. Deformation of a standard prosthesis when applying a mass of 90 kg

The greatest deformation is observed in the polyethylene insert, the point deformation zones are equal to a maximum of 7.47×10-3. Under the tibial augmentation, the cat deformation is equal to 5.76 x10-3, above the femoral augmentation, the deformation of the bone part is equal to 4.35 x10-3. To estimate the safety margin, the “Safety factor” function of the Simsolid program was used, according to which infographic images were obtained (Fig. 13).

Fig. 13. Safety margin of the prosthetic knee joint

The results obtained show that the lowest safety margin is for bone tissue, the highest for titanium components. In this case, the femur and tibia are red due to the negative impact of the prosthesis extensions on the entire studied part of the model. The effect of reaction contact forces and moments for a standard revision prosthesis is concentrated on the tibia augment. Under the influence of a mass of 90 kg, the displacement of this augment occurs downwards, the reaction forces are directed opposite to the action (Fig. 14).

Fig. 14. The action of reaction forces on implant components

In a standard prosthetic knee joint, there is friction occurring between the contact pair of the femoral component and the polyethylene insert. Under the action of the mass, the polyethylene insert creates a resistance reaction that is opposite to the action of the force acting on it, thereby playing a shock-absorbing role and reducing the dynamic load. Under the action of a force of 2.6 kN, the polyethylene insert creates a resistance reaction, the direction of which is opposite to the applied force. This mechanism provides a damping function, due to which the transmission of peak cyclic loads to bone structures and implanted elements is reduced. The calculated value of the total reaction force in the contact zone is 3.726×10³ N (3726 N). This corresponds to approximately 3.8 times the body weight of 90 kg, which corresponds to the biomechanical conditions of the stance phase during walking. In this phase of the walking cycle, the joint experiences the greatest dynamic pressure, which confirms the adequacy of the obtained simulation results to real biomechanical conditions. In addition, a moment of force of 27.6 N m was recorded in the model, which indicates the presence of a torque effect (torsional moment). Such a moment can lead to skewing or partial rotation of the implant, as well as to a non-uniform distribution of contact stresses between the femoral component and the polyethylene insert.

The moment of force M=27.6 N⋅m means that with a radius of the force arm of 0.04 m, the force is created:

F=M/r=27.6/0.04≈690 H

That is, the torque is equivalent to a lateral load of about 690 N applied at an angle to the axis of the knee. In clinical terms, this may contribute to increased wear of the polyethylene, instability of the endoprosthesis elements, or migration of the implant during prolonged use.

Thus, based on the obtained computer results, it becomes clear that only the use of individual extensions and augments minimizes the load on the residual bone. According to the conducted computer analysis, the advantage of using individual implants over standardized ones has been determined. When using an individual revision knee prosthesis with extensions of 120 mm, 130 mm and 150 mm, the most promising and least traumatic is the extension length of 130 mm. The weight acting on the joint with bone tissue defects creates deformation and displacement of the bone tissue, which is equal to 1.85 mm. This displacement is permissible, taking into account the elastic properties of the bone. With the use of this extension, in the prosthesis with individually designed components, the stress is within the permissible limits, which does not exceed the strength limit, namely, below 3.5 MPa.