Development and design of a carbon fiber insole intended for individuals with partial foot amputation Open Access

A person’s limbs, e.g., part of a foot, may be amputated due to an illness or an accident. In clinical practice, the ankle foot orthosis torque is modified, which is essential to account for each patient’s distinct body function and gait characteristics.¹ Partial amputation of the foot causes an uneven gait cycle, which is not energy efficient. Topping of the toe takes place during the swing phase. The possibility of foot drop has lessened with the development of prosthetic limbs, artificial body parts that amputees can use to replace missing body parts (limbs).² Prostheses are used to replace the amputee’s missing limbs, and they are made to be as comfortable and functional as the amputee’s original limbs. Improving the gait of patients with partial foot amputations requires the use of carbon fiber footplates or prosthetic sockets made using stereolithography or traditional techniques.^3,4 Since the normal gait pattern heavily depends on how much of the toe lever is still present after an amputation, the surgeon’s major objective is to remove as much of the foot as possible.⁵ To create an accommodating insole that is 3D printed, meets the clinical demands, and surpasses the current standard of care (SoC), clinician involvement is essential.⁶ The floor reaction force and ankle movement were taken into consideration when calculating the load conditions. This study investigates the idea that the main surgical objective required to preserve proper foot and ankle function should be to maintain the length of the residual foot.^7,8 The floor response force and ankle movement were considered while calculating the load conditions. People with rheumatoid arthritis (RA) are often prescribed foot orthoses as a form of treatment.^9,10 However, data on their cost-effectiveness under this condition are limited.¹¹ The author set out to evaluate two different types of foot orthoses for persons with established RA in terms of their clinical and financial efficacy.^12,13

The findings demonstrated that semi-rigid customized foot orthoses can reduce pain and disability scores when compared to standard insoles.^14,15 Despite this, the production of the unique foot orthoses was much more costly from a cost-effectiveness perspective, without resulting in a significant increase in cost per quality-adjusted life year (QALY).^16,17 The purpose of this paper is to estimate how well a specific orthotic stiffness tester performed while assessing the energy return and stiffness biomechanical characteristics of three distinct ankle–foot orthosis (AFO) designs.¹⁸ Using the orthotic stiffness tester, each AFO was put through several tests.¹⁹ Energy dissipation was assessed using the area inside the hysteresis loop, and stiffness was calculated using the average slope of the angle vs moment plot. While the tester was dependable for detecting stiffness, it was unreliable for evaluating energy storage.²⁰ 

This research outlines a systematic procedure for designing foot orthotics using simulation models validated through subsequent experiments. The current model has the potential to replace empirical tables in designing foot orthotics based on the weight and activities of end-users.²¹ It can be used to mimic many kinds of orthotics, loading scenarios, and material characteristics in order to get the desired results. In the end, the suggested model and its findings might serve as the basis for creating foot orthotics of the future. The approach involves systematically developing engineering analysis models using the finite element analysis (FEA) technique.²² The feasibility of employing additive manufacturing (AM) methods to create ankle–foot orthoses (AFOs) tailored to individual patients is tested. Using gait analysis data, the clinical performance of AFOs made from the PA12 material was evaluated in stroke survivors. An EinScan-Pro 3D scanner was utilized to capture the dimensions of the ankle and foot, and the AFO model was edited using the Geomagic Studio software. The initial AFO model was refined to the final design specifications, and the AFOs were then produced using multi-jet fusion (MJF) technology with the PA12 material.^23,24

Additive manufacturing (or 3D printing) enables rapid and cost-effective creation of innovative designs and objects using new materials. This technique, even with basic printers that may be used in offices, classrooms, workshops, and laboratories, is particularly beneficial for generating high-quality items with polymers and polymer composites. Many people can now readily test ideas or construct unique devices, thanks to it. It is easy to produce complex geometries that are difficult or impossible to achieve with existing technologies. Consequently, this technology has garnered interest across various fields, including the manufacturing of medical devices and prostheses, mechanical engineering, and basic sciences. It also allows for the use of materials that are difficult to machine traditionally. The authors demonstrate process advancements by detailing the production of printer parts designed to handle hard and abrasive polymer fillers. New nozzles with ruby inserts have been developed for such materials, proving durable and suitable for printing boron carbide composites.²⁵ Compared to traditional production techniques, additive manufacturing has several benefits, especially when large-scale customization is required.²⁶ The authors outline the procedure for producing a fully customized ankle–foot orthosis (AFO) using 3D printing. The patient’s foot geometry is first determined using photogrammetry. Subsequently, the data are converted and imported into the CAD modeling tool SolidWorksTM. The AFO is then parametrically modeled around the foot mesh to ensure that it can handle the anticipated mechanical stresses. The item is ultimately 3D printed using an FDM printer and tested on the patient.²⁷ 

Recently, additive manufacturing has emerged as a viable engineering approach for producing customized ankle–foot orthoses (AFOs). Compared to walking barefoot in people with drop feet, custom-molded PP-AFOs and customized SLS-AFOs have demonstrated significant improvements in spatiotemporal gait metrics and ankle kinematics.²⁸ On these gait and kinematic metrics, however, there was no statistically significant difference between the effects of PP-AFOs and SLS-AFOs.¹⁹ Foot drop can result from various diseases and injuries. The total incidence of foot drop has not been published, despite its frequency. The everyday activities of people affected by this illness are severely limited. Consequently, a systematic and optimal approach spanning multiple medical specializations is required for its diagnosis and treatment.²⁹ Disturbances at any point in the central or peripheral motor neuron circuit that supplies the foot’s dorsiflexor muscles, or at several points sequentially, can cause foot drop. Proper localization of the lesion(s) is essential for the right course of treatment and good results. The most frequent causes include damage to the peroneal nerve and L5 radiculopathy. When there is a reasonable likelihood of nerve recovery, a surgery performed by a neurosurgeon or spinal surgeon is a reasonable alternative. The authors think that surgical decompression at the fibular head, which entails little risk, should be made available to every patient with clinically diagnosed compressive neuropathy of the peroneal nerve and a severely painful foot drop.³⁰ It is recommended that individuals with Charcot–Marie–Tooth (CMT) disease use ankle–foot orthoses (AFOs). This study assessed how satisfied patients were with these orthotic treatments and devices, providing early information to enhance clinical care and the AFO design. Reduced satisfaction with pain, abrasions, and discomfort highlighted the need for improvements in AFOs. It will be essential to do more research in these areas to improve clinical care and provide guidance for the development of AFOs.^31,32 Individuals with lower limb amputations strive for enhanced functionality. Among active transtibial amputees, prosthetic feet designed for energy storage and return (ESR), typically made from carbon fiber composites, are the most popular. Recently, a new prosthetic foot made from fiberglass composites has been introduced to the market, but no comparative studies between these devices have been conducted. The findings of this study suggest that the newly introduced ESR foot made from fiberglass surpasses traditional carbon fiber designs.³³ 

The 3D template for this product is produced conventionally using the PFO production process. Figure 1 illustrates the process for creating orthoses using a 3D template. For the measurement of the foot, the UK 7 size foot is used. Epoxy resin is used to create layers after the carbon fiber is cut to UK 7 measurements. Finally, a load of 80 kg is applied to compress the fiber, unify it, and machine it to provide a smooth surface.

3D printed models are used to cut the product to the appropriate size and shape.³⁴ The present work employs the 3D printing technique to build a model before the actual manufacturing process starts, as depicted in Fig. 2. The 3D-printed model is used as the mold to produce carbon fiber footplates, and the model is drafted using Unigraphics-NX. The footplate design is imported as an STL file into the UP-mini 3.0 software, which prepares the model for 3D printing.³⁵ The printing process uses the ABS polymer, a durable and versatile plastic, to create the physical footplate with high precision, ensuring that it meets the required specifications.^23,36

Using the 3D-printed template, the carbon fiber sheets are cut according to the shape of the footplate, as shown in Fig. 3.

According to the foot measurement, carbon fibers sheets are cut. A footplate is manufactured with the layers of carbon fibers with the help of epoxy resin.^37,38 A load of 80 kg is applied so that the layers adhere together. After 30 min, the footplate is given the required shape. The carbon fiber footplate is manufactured as shown in the flow chart in Fig. 4.

Figure 5 depicts a partial foot orthosis crafted from carbon fibers. This lightweight, durable material provides robust support, enhancing mobility and comfort for individuals with partial foot amputations.^39,40

Figure 6 shows the length of foot, which is considered as a cantilever beam. The maximum deflection, which is referred to as deformation in theoretical calculations, is calculated for 20 and 40 kg loads at heel using the deformation formula,^41,42 where the span length L = 168.2 mm, the modulus of elasticity E = 290 GPa, and the moment of inertia I = 1470 mm⁴, as shown in Table I.

The deflection at toe is calculated theoretically for 20 and 40 kg loads, where the span length L = 178.4 mm, the modulus of elasticity E = 290 GPa, and the moment of inertia I = 3104 mm⁴, as shown in Table II.

To validate the footplate manufactured, a test rig was developed, where a mechanism is used to apply load using a load cell and measure the displacement. The test rig consists of a linkage arrangement for applying load at point 2 with the help of a digital vernier caliper and at point 1 where the toe or heel is fixed.^43,44

Figure 7 shows the simulated and actual test rigs of the current work. A partial foot orthosis is tested experimentally by using the test rig as shown in Fig. 7(a). To start, the load cell reading is reset to remove all previous loading information. Load is applied to the footplate at points 1 and 2 as shown in Fig. 7(b) to fix the toe and heel, respectively. Various loads are applied to appropriate areas before measuring the toe and heel’s deformation with a digital vernier caliper, and the results are noted down.

The carbon fiber footplate is tested at 20 and 40 kg loads, which are applied to the partial foot orthosis at the toe and heel. The toe is fixed, and the load is gradually applied to the heel. At 20 and 40 kg loads, deformation is measured at the heel using the difference between the initial and final readings of the vernier caliper to determine the deformation. The deformation at the heel is shown in Table III.

Loads of 20 and 40 kg are applied on the toe of the footplate during experimentation keeping the heel fixed. The deformation at the toe is measured using a digital vernier caliper and is calculated by the difference between the initial and final readings of the vernier caliper. Table IV tabulates the deformation at the toe at different loads.

Figure 8 demonstrates the application of a 20 kg (196 N) load at the heel, resulting in a deformation of 0.726 09 mm. These data provide insights into the material’s behavior under stress, showing how it responds to the applied force. The observed deformation highlights the mechanical properties and durability of the material, which is crucial for evaluating its suitability for structural and functional applications. The relationship between the load and deformation serves as an important parameter in understanding the material’s resilience and ability to maintain its integrity under varying levels of stress. These findings are valuable for optimizing the design and performance of the product.

The finite element analysis (FEA) results indicate that for both 20 and 40 kg loads applied at the heel, deformations of 0.726 09 and 1.4522 mm, respectively, were observed. These values are within the safe limits, as depicted in Fig. 9. The analysis demonstrates that the material can withstand the applied forces without exceeding critical deformation thresholds, ensuring the structural integrity and durability of the design. The proportional increase in deformation with the load confirms the material’s predictable response to stress, supporting its reliability in practical applications where such loads are encountered.

Figure 10 illustrates a deformation of 0.4385 mm under a 20 kg (196 N) load applied at the toe. This deformation provides valuable insights into the material’s response to stress specifically focused on the toe region. The relatively small displacement under this load suggests that the material has adequate stiffness and resistance when a force is applied in this area. Understanding the material’s deformation behavior in the toe region is crucial for ensuring its performance and durability, especially in applications where toe-directed stresses are common. These data support the material’s reliability and suitability for load-bearing tasks in similar scenarios.

Figure 11 depicts a deformation of 0.8188 mm under a 40 kg (392 N) load applied at the toe, indicating a substantial material response to the increased force. This deformation highlights the material’s behavior under significant stress, providing key insights into its structural performance under high-load conditions. The observed displacement reflects the material’s flexibility and its ability to absorb and redistribute force without failure. Such data are essential for assessing the material’s suitability in applications requiring durability and resistance to higher loads, ensuring that it can withstand real-world stressors while maintaining its integrity and functionality.

At a load of 20 kg, the theoretical deformation is 0.729 mm, the experimental value is 0.74 mm, and the analytical result is 0.726 09 mm. At a load of 40 kg, the theoretical deformation is 1.459 mm, the experimental value is 1.39 mm, and the analytical result is 1.4522 mm. The experimental values are slightly higher than the theoretical and analytical results at 20 kg, but they are slightly lower at 40 kg. The analytical values closely match the theoretical values, indicating that the analytical model accurately predicts the theoretical expectations.

At a load of 20 kg, the theoretical deformation is 0.4120 mm, the experimental value is 0.41 mm, and the analytical result is 0.438 mm. At a load of 40 kg, the theoretical deformation is 0.824 19 mm, the experimental value is 0.83 mm, and the analytical result is 0.818 mm. The experimental values are very close to the theoretical values, with minimal deviation. The analytical values are also close to the theoretical values, although there is a slight overestimation at 20 kg and a slight underestimation at 40 kg.

Tables V and VI show the comparative analyses of theoretical, experimental, and FEA results obtained from tests conducted at toe and heel.

The comparative analysis of the deformation results for the toe and heel regions at different loads provides valuable insights into the accuracy and reliability of theoretical, experimental, and analytical models. The following conclusions are made:

• At a load of 20 kg applied at the toe, the theoretical deformation is 0.729 mm, the experimental value is slightly higher at 0.74 mm, and the analytical result closely matches the theoretical result of 0.726 09 mm observed at heel. Similarly, at a load of 40 kg applied at the toe, the theoretical deformation is 1.459 mm, the experimental value of deformation is 1.39 mm, and the analytical value of deformation is 1.4522 mm.

• These observations indicate that the experimental values are slightly higher than both the theoretical and analytical results at 20 kg but slightly lower than those at 40 kg. The close match between the analytical and theoretical values suggests that the analytical model accurately predicts the theoretical expectations. Figure 12 shows the bar chart of the deformation at heel, and the percentage increase in deformation from 20 to 40 kg is ∼87.84%.

In the case of the heel region, at a load of 20 kg, the theoretical deformation is 0.4120 mm, the experimental value is very close at 0.41 mm, and the analytical result is slightly higher at 0.438 mm. At a load of 40 kg, the theoretical deformation is 0.824 19 mm, the experimental value is slightly higher at 0.83 mm, and the analytical result is very close at 0.818 mm. Figure 13 shows the bar chart of the deformation at the toe, and the percentage increase in deformation from 20 to 40 kg is ∼91.4%.

• Overall, the analytical model has shown a high degree of accuracy in predicting the theoretical deformations at both the toe and heel regions at different loads. The experimental values, while generally close to the theoretical values, exhibit some variation, which could be attributed to practical factors and measurement uncertainties.

By utilizing 3D printing and AI-driven modeling, personalized insoles tailored to each individual’s anatomy can be created, enhancing comfort and functionality, while exploring eco-friendly carbon fiber alternatives and recyclable materials can lead to more sustainable and environmentally conscious designs.

This work was supported by Researchers Supporting Project No. RSPD2024R990, King Saud University, Riyadh, Saudi Arabia.

The authors have no conflicts to disclose.

Chandrika Wagle: Data curation (equal). Prateek D. Malwe: Formal analysis (equal). Nitin P. Bhone: Investigation (equal). Naresh Jaiswal: Methodology (equal). Chetanraj D. Patil: Software (equal). Ahmed Fouly: Supervision (equal); Writing – review & editing (equal). Mohd Asif Shah: Conceptualization (equal).

The data that support the findings of this study are available from the corresponding author upon reasonable request.