Biodegradable hybrid polymer splint for tracheal regeneration and method of fabrication thereof

Abstract

The present invention generally relates to biomedical engineering and regenerative medicine. Specifically, the present invention relates to development of advanced tracheal splints for the treatment of tracheal disorders. More particularly, the present invention relates to a process for preparation of biodegradable hybrid polymer splints, the process comprising, a) carrying polymerization reaction of polycaprolactone (PCL) with pectin in presence of a catalyst to obtain pectin-g-PCL; and b) 3D printing of pectin-g-PCL. The present invention also relates to a biodegradable hybrid polymer splint comprising honeycomb structure, rectangular convolutions or combination thereof. Further, the tracheal splints are designed to support tracheal regeneration, enhance cellular interaction, and eliminate the need for additional surgical procedures for removal.

Description

[0001] BIODEGRADABLE HYBRID POLYMER SPLINT FOR TRACHEAL

[0002] REGENERATION AND METHOD OF FABRICATION THEREOF

[0003] TECHNICAL FIELD OF THE INVENTION

[0004] The present invention generally relates to biomedical engineering and regenerative medicine. The present invention relates to development of advanced tracheal splints for the treatment of tracheal disorders. Specifically, the present invention relates to a process for preparation of biodegradable hybrid polymer splint. Further, the tracheal splints are designed to support tracheal regeneration, enhance cellular interaction, and eliminate the need for additional surgical procedures for removal.

[0005] BACKGROUND OF THE INVENTION

[0006] In the field of medical progress, the combination of 3D printing technology with biodegradable materials has brought about a new wave of innovation, particularly evident in the creation of tracheal splints. Tracheal disorders encompass a wide range of conditions affecting the structure and function of the trachea. One common example is tracheomalacia, a condition characterized by weak cartilage in the trachea that leads to the collapse or narrowing of the airway during breathing. Another condition, tracheal stenosis, involves the narrowing of the tracheal lumen due to scarring or inflammation, leading to difficulty in breathing. Surgical techniques for tracheal disorders can be effective however, surgical techniques have inherent disadvantages.

[0007] Tracheal reconstruction and resection may lead to complications such as infection and anastomotic issues. Tracheostomy can result in infections and scarring. Treatment options, such as stenting the trachea with non-biodegradable materials like silicone or metal, are effective in providing immediate support to the airway. However, these materials may lead to long-term complications such as inflammation, tissue damage, or the need for additional surgeries to remove the stent. The emergence of biodegradable tracheal splints offers a promising alternative for the management of tracheal disorders. These splints are designed specifically for each patient using their own imaging data, ensuring a perfect fit to their unique tracheal anatomy. This customized support helps to strengthen the weakened trachea, maintaining an open airway and enhancing respiratory function. Additionally, biodegradable materials used in these splints have several advantages over traditional non-biodegradable options. As splint slowly degrades within the body over time, there is no need for additional surgeries to remove it once it has served its purpose. This reduces the risk of long-term complications and streamlines the overall treatment process for patients with tracheal disorders. In the field of tracheal tissue engineering, researchers have been investigating various polymers to develop innovative solutions for tracheal repair and regeneration. Among the polymers studied, polylactic acid (PLA) and its copolymer poly(lactic-co-glycolic acid) (PLGA) have received significant attention due to their biocompatibility, biodegradability, and adjustable mechanical properties. These polymers offer flexibility in fabrication methods such as in electrospinning and 3D printing, which enable creation of scaffolds with customized structures and degradation rates. Polycaprolactone (PCL) has also been explored for its favourable mechanical properties and slower degradation, making it suitable for long-term support in tracheal tissue engineering applications. Other polymers like polyethylene glycol (PEG) and polyurethane (PU), have also been studied for hydrophilicity, flexibility, and biocompatibility. By utilizing these polymers and their intrinsic properties, researchers strive to develop biomimetic scaffolds that facilitate tracheal tissue regeneration and ultimately enhance outcomes for patients with tracheal disorders.

[0008] Proteins and other natural polymers, such as polysaccharides, offer an advantage over synthetic polymers because they are more similar to living tissue. Tracheal tissue engineering extensively explores natural polymers such as proteins and polysaccharides, including chitosan, gelatin, silk fibroin, collagen, hyaluronic acid, and fibrin. These natural polymers are studied in various forms such as hydrogels, foams, coatings, and electrospun fibers. However, while these natural polymers offer biocompatibility and support cellular growth, they alone often lack the mechanical strength required to maintain the patency of the airway.

[0009] Among natural polymers, pectin remains unexplored in the field of tracheal tissue engineering. Pectin is a preferred material for the building of scaffolds for tissue engineering applications because it shares structural similarities with the glucosaminoglycans (GAG) in the extracellular matrix of human cells, which promote inter- and intracellular connections. Currently, a wide variety of pectin-based scaffolds are used for this purpose, including sponge, electrospun fibre, hydrogel, 3D printed structures, and others. Due to a lack of bioactive functional groups and its inherent hydrophobicity, PCL’s ability to promote cell adhesion and proliferation is severely constrained. The limitations caused by the surface features of PCL -based scaffolds have led to the development of numerous modifications, such as coating or blending of PCL with other biocompatible materials in order to generate biocompatible scaffolds with improved mechanical strength and inherent bioactivity for tissue engineering. However, challenges include the potential compromise of mechanical properties, variations in degradation rates leading to uneven tissue breakdown, and processing complexities to achieve uniform distribution.

[0010] Frejo, L. et al, (Bioelectron Med 5, 15 (2019)) report overview of 3D-bioprinted tracheal reconstruction and how PCL has become the most used biomaterial in 3D-printing in designing customized airway stents. Morrison, Robert J., et al. (Science translational medicine 7.285 (2015): 285ra64-285ra64) reports mitigation of tracheobronchomalacia with 3D-printed personalized medical devices in paediatric patients. However, these materials are not stable and require sufficient amount of solvents for 3D printing. Additionally, commercially available polycaprolactone, which has a longer degradation period.

[0011] Inherent polymer chain modification is a more efficient technique compared to surface modification. In mass modification, cell signaling peptides are integrated directly into the biomaterials, ensuring that recognition sites are present throughout the material, both within its mass and on its surface. Efforts have been dedicated to achieving the necessary mechanical properties and tissue regeneration in tracheal tissue engineering by exploring diverse combinations of natural and synthetic polymers. However, the inherent modification of PCL with pectin within the realm of tracheal tissue engineering remains unexplored.

[0012] Therefore, there is a need for addressing these gaps in developing innovative tracheal substitutes and include pectin into PCL to improve the properties of PCL.

[0013] OBJECTIVE OF THE INVENTION

[0014] An objective of present invention is to synthesize pectin-grafted PCL using optimized polymeric reaction conditions, ensuring the production of high-quality material for tracheal splint fabrication. Another object of the present invention is to design tracheal splints that are biodegradable and bioresorbable, eliminating the need for additional surgical procedures for removal.

[0015] Another object of the present invention is to create tracheal splints using 3D printing technology, which offer enhanced bio-interaction and support tracheal regeneration.

[0016] SUMMARY OF THE INVENTION The present invention generally relates to biomedical engineering and regenerative medicine. Specifically, the present invention relates to development of advanced tracheal splints for the treatment of tracheal disorders. More particularly, the present invention relates to process for preparation of a biodegradable hybrid polymer splint. The tracheal splints are designed to support tracheal regeneration, enhance cellular interaction, and eliminate the need for additional surgical procedures for removal.

[0017] In an aspect, the present invention relates to a process for preparation of a biodegradable hybrid polymer splint, the process comprising a) carrying polymerization reaction of polycaprolactone (PCL) with pectin in presence of a catalyst to obtain pectin-g-PCL; and b) 3D printing of pectin- g-PCL to obtain the biodegradable hybrid polymer splint.

[0018] In another aspect, the present invention relates to a biodegradable hybrid polymer splint comprising honeycomb structure.

[0019] BRIEF DESCRIPTION OF THE DRAWINGS

[0020] Figure 1 illustrates1H NMR spectra of Pec-g-PCL, PCL and Pectin of Example 1.

[0021] Figure 2 illustrates FTIR Spectrum of Pec-g-PCL, PCL and Pectin of Example 1.

[0022] Figure 3 illustrates thermal analysis of pectin-g-PCL a, b) TGA and DTGA curves of pectin-g- PCL with various feed ratios c, d) DSC analysis of pectin-g-PCL e)XRD analysis of Pec-g-PCL f) contact angle measurements of Pec-g-PCL of Example 1.

[0023] Figure 4 illustrates rheological testing of Pec-g-PCL polymers of Example 1.

[0024] Figure 5 illustrates 3D printed structures of Pec-g-PCL of Example 3.

[0025] Figure 6 illustrates design and fabrication of tracheal splint (Model 1-M1 and Model 2-M2) of Example 5.

[0026] Figure 7 illustrates mechanical study of 3D printed splints a, b, c) Radial compression of M1 and M2 d, e, f) Lateral compression of M1 and M2 f) Cyclic radial compression of M1 and M2 g, h) cyclic compression of M1 and M2 of Example 5.

[0027] Figure 8 illustrates finite elemental analysis of tracheal splints (M1 and M2) of Example 5, a-f) Radial compression g-1) Lateral compression.

[0028] Figure 9 illustrates degradation studies of Pec-g-PCL scaffolds a-c) Accelerated degradation d-f) Physiological degradation studies of cylindrical scaffold (5 mm x 5 mm x 6 mm ((width x diameter x height)). Figure 10 illustrates SEM morphological study of accelerated and physiological degradation of commercial PCL and Pec-g-PCL scaffold (Model 2 of Example 5).

[0029] Figure 11 illustrates in-vitro biocompatibility test of Pec-g-PCL scaffold (Model 2 of Example 5). Figure 12 illustrates the scheme for synthesis of Pec-g-PCL via ring opening polymerization.

[0030] DETAILED DESCRIPTION OF THE INVENTION

[0031] The present invention generally relates to biomedical engineering and regenerative medicine. Specifically, the present invention relates to development of advanced tracheal splints for the treatment of tracheal disorders. More particularly, the present invention relates to a process for preparation of biodegradable hybrid polymer splint. The tracheal splints are designed to support tracheal regeneration, enhance cellular interaction, and eliminate the need for additional surgical procedures for removal.

[0032] In an embodiment, the present invention provides a process for preparation of a biodegradable hybrid polymer splint, the process comprising: a) carrying polymerization reaction of polycaprolactone (PCL) with pectin in presence of a catalyst to obtain a pectin-g-PCL; and b) 3D printing of pectin-g-PCL of step (a) to obtain the biodegradable hybrid polymer splint.

[0033] The PCL used in step a) is obtained from polymerization of ε-caprolactone (ε-CL) monomer. Step a) is also referred to as graft polymerization process. The polymerization process is initiated by the hydroxyl groups of pectin and catalyzed by the catalyst. In step a), the ratio of pectin to PCL (polycaprolactone) is in a range from 1:5 to 1:1000. The catalyst is selected from a tin-based catalyst and is present in an amount ranging from 0.1 mol% to 1 mol%. Preferably, the catalyst is stannous octoate (Sn (Oct)2). Preferably, the step a) is carried out at a temperature in a range from 100 °C to 200°C for a period ranging from 16 hours to 48 hours. Preferably, step a) is carried out in an inert atmosphere of nitrogen, argon or mixture thereof.

[0034] The step a) is carried out by bulk ring opening polymerization (ROP) to obtain pectin-g-PCL. The synthesis of Pec-g-PCL via ring opening polymerization is shown in Figure 12.

[0035] Then, in step b) the 3D printing of pectin-g-PCL is carried out by hot melt extrusion. The step b) is carried out at a temperature in a range from 70°C to 100°C to obtain the hybrid polymer splint. The process of the present invention is a one -pot, solvent-free process that enables the polymerization of natural polymer pectin with the synthetic biodegradable polymer (PCL), thereby providing improved mechanical strength and biocompatibility of the splints.

[0036] The step of 3D printing ensures consistent layer-by-layer adhesion, resulting in reproducible mechanical and biological properties and facilitates the creation of hybrid polymer splint (scaffolds) with controlled porosity and pore size, thereby enhancing nutrient diffusion, cell infiltration, and vascularization, which are critical for tissue regeneration. Additionally, the step of 3D printing enables the fabrication of splints comprising honeycomb structure, which can be tailored to mimic the natural architecture of the trachea and optimize mechanical performance.

[0037] In another embodiment, the present invention provides a biodegradable hybrid polymer splint comprising honeycomb structure, rectangular convolutions or combination thereof. The honeycomb structure may also be referred to as hexagonal structures. The honeycomb structure consists of a hollow bellows scaffold meticulously designed to mimic the natural trachea’s architecture.

[0038] The biodegradable hybrid polymer splint of the present invention has inner diameter ranging from 5 mm to 15 mm, a length ranging from 10 mm to 20 mm, and a wall thickness ranging from 1 mm to 3 mm.

[0039] In an embodiment, the biodegradable hybrid polymer splint comprising rectangular convolution features a longitudinal thickness of 1.5 mm per convolution, with a separation distance of 2 mm, a luminal diameter of the bellows at 5 mm, an inner diameter of 10 mm, a length of 15 mm, a wall thickness of 3 mm, and suture holes spaced at 2 mm intervals.

[0040] In another embodiment, the biodegradable hybrid polymer splint comprising honeycomb structure features an inner surface of the lumen patterned with equilateral hexagons, each having an interior angle of 120 degrees. Each hexagon is marked with 2 mm diameter circles to create porous splints optimizing airflow and tissue integration. The lumen surface further comprises double convoluted rectangles with thicknesses of 2 mm and 1 mm. The tracheal splints have overall dimensions of 12 mm in height, 12 mm in length, and 12 mm in width. The biodegradable polymer splint of the present invention has a porosity of at least 70%.

[0041] Further, the biodegradable hybrid polymer splint of the present invention, in in-vitro cell culture assays demonstrate cytocompatibility with tunable mechanical properties tailored to different geometries. The degradation rate of biodegradable hybrid polymer splint of the present invention exceeds that of commercial PCL splints, with enhanced cell proliferation, indicating its potential for tissue regeneration applications. Additionally, the biodegradable hybrid polymer splint exhibits mechanical properties similar to native tracheas thereby suggesting its utilization for treating tracheal disorders, potentially extending patient life. The biodegradable hybrid polymer splint incorporates a fixed open angle of 90° to ensure proper positioning of the splint over the airway.

[0042] In certain embodiments, the biodegradable hybrid polymer splint of the present invention is bioresorbable which eliminates the need for additional surgical procedures for removal of the said splint and further due to its biodegradability property is environmentally friendly. In certain embodiments, the biodegradable hybrid polymer splint exhibits 30% of degradation over 35 days.

[0043] EXAMPLES

[0044] The present invention is further explained in the form of the following examples. However, it is to be understood that the following examples are merely illustrative and are not to be taken as limitations upon the scope of the invention.

[0045] Materials

[0046] Low methoxyl pectin from citrus peel (galacturonic acid ≥74.0% on dried basis) from Sigma- Aldrich. The degree of esterification of pectin was 60% with an average molecular weight of 30000-100000 g mol-1. ε-caprolactone (ε-CL) from Sigma-Aldrich (>99%) was dried over calcium hydride for 24 hours and distilled under reduced pressure. Stannous octoate from Sigma Aldrich is used as a catalyst. All other reagents were used as such unless it is specified.

[0047] Example 1: synthesis of Pec-g-PCL via ring opening polymerization

[0048] Bulk ring opening polymerization (ROP) was carried out to synthesize Pec-g-PCL as per Figure 12. Various compositions of pectin and caprolactone monomer, ranging from ratios of 1 :5 to 1 : 100, were prepared. Requisite amounts of Pectin (0.05g- 1g) were weighed and transferred to a 50mL round bottom flask equipped with Teflon coated stir bar under an inert N2atmosphere. Then 5 g of ε-caprolactone (ε-CL) monomer was charged in a round bottom flask in a stepwise fashion with 0.4 gram of stannous octoate as catalyst. Polymeric reaction conditions were optimized with respect to time, temperature, and compositions comprising different ratios of pectin to PCL. The polymers comprising different ratios of Pec:PCL along with the process parameters employed are presented in Table 1 below. The reaction mixture or polymer compositons then obtained was solubilized in acetone and precipitated in cold water. The precipitated pectin -g-PCL compound was centrifuged at 10,000 rotation per minute (RPM) for 10 min and dried under vacuum for 4 h. Table 1:

[0049] Example 2: Characterization of Sample 1 to 5 prepared according to Example 1

[0050] Pec-g-PCL samples’ chemical shifts were analyzed using a 400 MHz1H- Nuclear Magnetic Resonance (NMR) Bruker Spectrometer, employing CDCI3 as the solvent. FTIR spectroscopic analysis was conducted using a PerkinElmer Fourier-transform infrared (FTIR) spectrometer at room temperature. FTIR-ATR spectra were obtained by averaging 20 scans within the 400-4000 cm-1range. To assess the crystallinity of the polymers, XRD analysis was performed using an XRD machine (Rigaku, Tokyo, Japan). X-ray intensity was measured within the 20 range of 10 to 30°, utilizing CuKα radiation with a wavelength of 1.54 A and a scan rate of 0.05 s-1. The degree of crystallinity (Xc) was determined using: where Acrepresents the crystalline area and Aarepresents the amorphous area. Additionally, contact angle measurements were obtained using a goniometer system equipped with a digital camera, employing the sessile drop method. A drop of distilled water (30 μl) was carefully placed on the polymer film, and the mean contact angle was computed as the average angle within 2 to 5 seconds after the droplet was deposited onto the surface.

[0051] Figure 1 represents the1H- NMR spectra of Pec-g-PCL prepared according to Example 1 (Sample nos. 1 to 5) and compared with Pectin and PCL individually. The said figure provides a comparative analysis of the chemical shifts and peak intensities corresponding to the structural components of pectin, PCL, and their grafted copolymers prepared according to Example 1. PCL shows CH2peaks of the opened ring caprolactone monomer at δ=1.32 ppm, 1.58 ppm, 1.65 ppm, 2.32 ppm, 3.61 ppm. Peak at δ= 4.06 ppm corresponds to CH2bound to oxygen atom in PCL backbone. Pectin shows its characteristic CH peak at δ=3.39 ppm, 3.74 ppm, 3.77 ppm, 4.13 ppm, 4.38 ppm, 4.58 ppm, 4.70 ppm and 5.34 ppm. The hydroxyl group (OH) on the backbone of pectin was observed at 3.59 ppm. It is essential to note the peaks corresponding to pectin and PCL displayed minor shifts, owing to chemical interactions in the pectin-g-PCL sample. The1H-NMR spectra confirmed the effective ROP and grafting of ε-CL to the surface of the pectin polysaccharide.

[0052] Figure 2 depicts the FTIR spectra of Pec-g-PCL prepared according to Example 1, pectin and PCL. The distinctive peak detected in the region 3462 cm-1for pectin corresponds to secondary alcohol OH stretching. Symmetric and asymmetric C-H stretching was observed at 2954cm-1. C=O characteristic peaks are detected in the range 1750-1700cm-1. The typical peak developed in the 1170-1150cm-1range, which is related to C-O-C stretching vibration. The C=O stretching of PCL and pectin overlapping in the range 1750-1700cm-1, forming a high-intensity peak. The typical pectin and PCL peaks were observed in the Pec-g-PCL which confirms the successful grafting of PCL on pectin.

[0053] Thermal analysis

[0054] The thermal stability of Pec-g-PCL polymers prepared according to Example 1 was assessed using a Thermogravimetric Analyzer (PerkinElmer, Pyris Diamond System). Samples 1 to 5 prepared according to Example 1 weighing approximately 10 mg underwent thermal ramp from 50 to 900°C at a heating rate of 10°C / min, under a constant nitrogen flow of 50 ml / min. Differential Scanning Calorimetry (DSC) (TA Instruments DSC250) was employed to investigate the thermal behaviour of samples 1 to 5 of pectin-g-PCL polymers prepared according to Example 1. For non-isothermal DSC analysis, each sample of Example 1 weighing between 9 and 10 mg were heated from -65 to 100°C at a rate of 10°C / min, held for 5 minutes to erase the thermal history, cooled to -65°C at the same rate, and subsequently reheated to 100°C. All experiments were conducted under nitrogen atmosphere (flow rate: 50 ml / min) to prevent oxidative degradation. The degree of crystallization (Xc) of the polymers was determined using Equation where ΔHm represents the enthalpy of fusion of pectin-g-PCL samples prepared according to Example 1 (in J g-1), and AHioo% represents 100% crystallinity (139.5 J g-1) The enthalpies of fusion were derived from the area under the endotherm curve.

[0055] Figure 3 a depicts the thermal analysis of pectin-grafted polycaprolactone (pectin-g-PCL) polymers, by thermogravimetric analysis (TGA); Figure 3b depicts the derivative thermogravimetric analysis (DTGA); Figure 3c and 3d depicts the differential scanning calorimetry (DSC), Figure 3e depicts X-ray diffraction (XRD), and Figure 3f depicts contact angle measurements. The thermal behavior of pectin-g-PCL was investigated by means of TGA under nitrogen atmosphere.

[0056] Figure 3 a and 3b shows characteristic TGA and derivative TGA curves of PCL and pectin-g-PCL. It was observed that pectin-g-PCL of different feed ratios (Samples 1 to 5 of Example 1) exhibits a single degradation stage with initial degradation around 270-305 °C at Tio (10%) weight loss and Tmaxof around 348-412 °C leading to total decomposition of the polymeric chain. The DTGA curve in Figure 3b depicts the maximum temperature for thermal degradation of pectin-g-PCL with different feed ratios. The TGA results (Figure 3a and 3b) suggests that the thermal properties of pectin-g-PCL prepared according to Example 1 was not found to be affected by the addition of PCL chain into pectin, instead it improved the thermal stability of grafted polymers.

[0057] DSC was used to determine the melting point and crystallinity of Pec-g-PCL as represented in Figure 3c and 3d. The DSC data showed that when PCL concentrations increased, the enthalpy of fusion decreased, which significantly impacted the degree of crystallinity. The melting enthalpy was computed by integrating the melting endotherm and dividing that amount by the enthalpy of fusion of totally crystalline PCL, which is 139.5 J / g in the literature, to get the degree of crystallinity (Xc%) of samples. The assessment of the degree of crystallization (i.e., Xc (%)) from the DSC thermogram (Figures 3c and 3d) in depicted in Table 2 below. Table 2 suggests that as the concentration of caprolactone (PCL) increases, the degree of crystallization rises considerably, thereby leading to a polymer (pectin-g-PCL) of well-aligned structure. Figure 3e depicts the X-ray Diffractogram (XRD) of pure pectin and pectin-g-PCL of Sample 1 to 5 of Example 1. The diffractogram of the pure pectin showed two broad halos at 12.9° and 22° of 29, which is in line with what has been reported in the literature for pure pectin films. In the spectrum of Pec-g-PCL, the typical PCL peaks at 20 = 21.5° and 24.5°, which is attributed to semicrystalline diffraction of (110) and (200) lattice planes due to the well-known orthorhombic crystal lattice of PCL indicating lower crystallinity compared to pure PCL samples were observed. The degree of crystallinity was calculated according to the following equation (1)

[0058] -Eq 2where Acis crystalline area and Aais the amorphous area. The crystallinity calculated in presented in Table 2 below. Figure 3e and Table 2 suggests that as the concentration of pectin increases, the percentage of crystallinity decreases, as expected. This decrease is due to the presence of amorphous pectin chains, which, while hidden by the crystallinity of PCL, hinder the crystallization of PCL. The decrease in crystallinity observed in Pec-g-PCL samples of Example 1 is promising for our current approach, suggesting that a lower crystalline structure facilitates favorable conditions for cell adhesion and tissue growth.

[0059] Table 2: Thermal properties of Pec-g-PCL

[0060] Figure 3f depicts the contact angle measurements assessing the hydrophilicity and hydrophobicity of pectin-g-PCL polymer samples prepared according to Example 1. Pectin is hydrophilic in nature and PCL is hydrophobic in nature. The contact angle increases from Pec-g-PCL (1:5) to Pec-g- PCL (1: 100) as the concentration of CL increases which indicates the gradual increase in hydrophobicity of the material.

[0061] Overall, FIG. 3 demonstrates that the thermal properties, crystallinity, and surface characteristics of pectin-g-PCL samples 1 to 5 prepared according to Example 1 can be tuned by adjusting the polymer composition, which is important for optimizing the material for preparation of biodegradable hybrid tracheal splints.

[0062] Rheological testing of pectin-g-PCL samples (1 to 5) prepared according to Example 1

[0063] Figure 4 depicts the rheological testing of pectin-grafted polycaprolactone (pectin-g-PCL) polymers of the present invention with varying pectin-to-PCL ratios, evaluating their suitability for 3D printing applications.

[0064] The rheological properties of the synthesized polymers of Example 1 were assessed using parallel plate geometry. Hot-pressed polymer discs with a diameter of 25 mm and a thickness of 2 mm were prepared for the study. An amplitude sweep experiment was conducted at 70°C, employing a frequency sweep of 1 rad / s to determine the linear viscoelastic region (LVR). Subsequently, oscillatory experiments were performed using a strain percentage from the LVR region. The polymeric samples were subjected to angular frequencies ranging from 0.1 to 100 rad / s at 70°C, with a constant strain value of 1%. To quantify the shear-thinning behavior of the polymers, the power law model ( ) was utilized. In oscillatory temperature sweep mode, a temperature ramp from 60 to 140°C at a heating rate of 5°C / min and a strain rate of 1 % were applied to evaluate the dynamic complex modulus and viscosity. Additionally, the isothermal stability of the material at 80°C was assessed through time sweep experiments. Prior to the initiation of each experiment, all samples were equilibrated for 2 minutes to ensure a uniform temperature throughout the sample. According to rheological theory, optimal printing inks for 3D printing should demonstrate viscoelastic properties. This entails the ability to extrude as a viscoelastic fluid, displaying shear- thinning behavior under shear stress. Furthermore, they should transition into a self-supporting viscoelastic solid post-extrusion, with adequate viscosity to maintain structural stability. To study the rheological characteristics and assess the printability of materials, the Pec-g-PCL ( 1 :5 to 1 : 100) of Example 1 was assessed for their elastic and viscous nature by evaluating the storage modulus (G') and the loss modulus (G") to understand the extrudability of the Samples 1 to 5 of Example 1 at 70 °C as depicted in Figure 4a and 4b. Figure 4a included amplitude sweep experiment and Figure 4b included frequency sweep experiment. Before conducting oscillatory measurements, the LVR region is identified to isolate regimes where the material structure remains intact.

[0065] In the amplitude sweep findings, the said Figure 4a, demonstrated that the storage modulus (G') and loss modulus (G”) remained nearly constant, indicating that the polymer structure was unaffected. In the strain sweep experiment the pectin-g-PCL samples of Example 1 having ratio of pectin : PCL of 1:5 and 1 :10 shows a crossover of G’ and G” at a higher strain rate. The rheological behaviour of the pectin-g-PCL exhibits notable variations depending on the percentage of pectin incorporated. When a higher proportion of pectin is present, the Pec-g-PCL tends to display a transition from predominantly elastic to viscous behaviour as strain rates increase. Initially, the storage modulus (G) surpasses the loss modulus (G"), indicating a more elastic response. However, with increasing strain rates, the viscous nature of Pec-g-PCL becomes increasingly prominent, causing a crossover where the loss modulus exceeds the storage modulus. This shift underscores the material’s transition from predominantly elastic deformation to more viscous flow. While for pectin-g-PCL samples 3 to 5 having ratio of pectimPCL of 1:25 to 1: 100 the G’ and G” are nearly constant. When a higher percentage of PCL is incorporated into the pectin-g-PCL polymer, the rheological behavior typically shows that the loss modulus (G") surpasses the storage modulus (G'). This means that the material becomes more viscous than elastic under the applied strain conditions, despite PCL’s normally rigid and crystalline nature. Overall, the viscoelastic behaviour of the composite shifts towards a predominantly viscous response. As a result, a 1 % strain (within the LVR) was used for the remaining experiments.

[0066] Further, as observed from the frequency range tests as shown in Figure 4b, the storage modulus (G') of all polymers was less than the loss modulus (G"), indicating the pseudo-plastic property (shear-thinning behaviour) and the extrudability of the pectin-g-PCL samples of Example 1 . As, in 3D printing, it is important to maintain continuous ink flow, which depends on the shear- thinning behavior of the printing inks once the ink flow begins. The pseudo-plastic property and the extrudability of the pectin-g-PCL samples confirms their suitablility for 3D printing as the said properities enable the easy extrusion of the sample from the nozzle during the process of 3D printing. Further, the shear-thinning behaviour enables the controlled deposition on the material during the process of 3D printing.

[0067] Figure 4c shows the complex viscosity curve of Pec-g-PCL samples of Example 1 over a temperature range of 60-140°C. For polymer compositions of 1 :5 and 1: 10, the complex viscosity remains relatively constant at elevated temperatures, which makes extrusion challenging due to the lack of shear-thinning behavior. On the other hand, for ratios of 1 :25 to 1 : 100, the viscosity of polymer melts initially decreases sharply with increasing temperatures, followed by a gradual reduction in the rate of viscosity decrease. From the said Figure, an optimal temperature of 70°C was identified as a printing temperature to facilitate extrusion, as this temperature benefits from decreased viscosity and the onset of shear-thinning behavior.

[0068] Figure 4d demonstrates the thermal stability of the pectin-g-PCL polymer samples of Example 1, tested through dynamic time sweep experiments at a constant frequency of 10 rad / s and a temperature of 70°C. The consistency in modulus (G and G”) over time indicates no degradation. Further, the said Figure demonstrates that the polymer sample 1 to 5 of Example 1 demonstrated sufficient stability for up to 30 minutes, allowing enough time for 3D printing experiments to be completed. Typically, a scaffold of dimensions 12mm x 10mm x 10mm can be completed within 20 minutes.

[0069] Example 3: 3D printing of polymer samples (1 to 5) obtained in Example 1

[0070] To initiate the printing process, polymers were loaded into the nozzle of a high-temperature dispensing head and dispensed through the cylindrical metal needle with an inner diameter of 27 gauge using pneumatic pressure, crosshead speed, at a temperature of 70°C for 10 minutes, reaching a molten state for extrusion. The process parameters are depicted in Table 3 below. The 3D printing of each polymer sample took approximately 30 minutes to complete.

[0071] Table 3 :

[0072] Printing parameters for 3D printing Pec-g-PCL

[0073] Example 4: Micro-computed tomography (μ-CT) imaging of 3D printed polymer samples obtained in Example 3

[0074] The polymer samples obtained in Example 3 were analysed by X-ray tomographic imaging using a Bruker Skyscan 1275 Micro CT, @30 kV with no filter, 360° and 2 frames per second scanning rate. The procedure began with the acquisition of scan data, followed by reconstruction utilizing Sky Scan NRecon GPU software. Representative 2D projection and reconstructed images were generated by applying a threshold between the two local maxima in the grey-scale histogram. For optimal alignment, datasets underwent potential reorientation using SkyScan Dataviewer software. Subsequently, the data analysis occurred in both 2D and 3D using SkyScan CTAn software. To visualize the structures comprehensively, volumetric 3D models were crafted utilizing SkyScan CT Vox software as depicted in Figure 5.

[0075] Figure 5 depicts the 3D printed structures of pectin-grafted polycaprolactone (pectin-g-PCL) of Example 3. The said figure highlights the effect of pectin concentration in Pec-g-PCL sample on the fabrication of 3D printed structures. As observed from the said Figure, the Pec-g-PCL with a lower CL monomer concentration was difficult to print as the viscosity of the polymer was lower compared to the composition with a high CL concentration.

[0076] As the flowability of polymer plays a vital role in additive manufacturing (3D printing), amongst the samples of Example 3 comprising different ratios of pectin: PCL, the polymer Sample no. 3 comprising pectin: PCL ratio of 1:25 showed the best hydrophilicity, rheological behaviour, and printability. The said polymer sample 3 was further used to build biodegradable hybrid polymer splints. Example 5: Process for preparation of biodegradable hybrid polymer splints

[0077] To understand how the design and porosity of the material affect mechanical properties and tracheal tissue regeneration, two different models were tested and compared. The tracheal splint’s unique designs were created using two different CAD modeling software programs: CATIA V5 R20 and FreeCAD. The tissue-engineered biodegradable hybrid polymer splint structure consisted of a honeycomb structure comprising hollow bellows scaffold meticulously designed to mimic the natural trachea's architecture. The initial concept featured an open, bellows-shaped cylinder with rectangular convolutions strategically placed to accommodate the placement of the splint outside the collapsed airway. The designs incorporated a fixed open angle of 90° to ensure proper positioning of the device over the airway. For the first model (Model 1, M1), each convolution was designed with a longitudinal thickness of 1.5 mm and a separation distance of 2 mm with the luminal diameter of the bellows scaffold set at 5 mm. Key specifications included an inner diameter of 10 mm, a length of 15 mm, a wall thickness of 3 mm, and suture holes spaced at intervals of 2 mm. These parameters were carefully chosen to optimize the functionality and compatibility of the tracheal splint. The second model (Model 2, M2), was designed, inspired from the composite structure of the trachea, which contains both rigid and flexible elements. To mimic this, the inner surface of the lumen of tracheal splint was patterned with equilateral hexagons, each with an interior angle of 120 degrees. These hexagons were marked with 2 mm diameter circles, which helped create porous splints with optimized airflow and tissue integration. Additionally, the lumen surface was separated by double convoluted rectangles with thicknesses of 2 mm and 1 mm. The hexagon pattern (honeycomb structure) was considered as an approach to optimize airflow and tissue integration. The incorporation of double convoluted rectangles between these hexagonal patterns added an interesting structural element, likely providing both flexibility and stability to the overall design. This combination of geometric shapes seems well-suited for mimicking the natural properties of the trachea while ensuring functionality and compatibility with surrounding tissues. Both models had dimensions of lOmmxlOmmxlOmm obtained using an Allevi 3 bioprinter model from Allevi Inc. in Philadelphia, PA, USA, equipped with hot melt extrusion capabilities ranging from 20 to 160°C. The printing parameters included a temperature of 70°C, printing pressure of 10 Psi, nozzle gauge size of 30, layer height of 0.2 mm, printing speed of 1 mm / s, and print bed temperature of 25°C. The predetermined designs of the tracheal splints were uploaded as STL files to the Bioplotter. Example 6: Evaluation of the splint of Example 5

[0078] In tissue engineering, the relationship between cells and their surroundings is closely linked to the structure of the scaffold (splint). This is because the extracellular matrix helps facilitate specific interactions between integrins and ligands, which can affect cell proliferation. Therefore, the 3D scaffold environment must have certain characteristics such as high porosity, appropriate pore size, and an interconnected pore network. These features are crucial for promoting optimal cell growth, migration, attachment, and proliferation. Thus, the splint obtained in Example 5 was evaluated using Micro-Computed Tomography (Micro-CT) imaging to assess these structural characteristics. Micro-Computed Tomography (Micro-CT) imaging

[0079] Micro CT imaging (Figure 6) was used to create detailed three-dimensional visualizations of the 3D printed Pec-g-PCL tracheal splints of Example 5, allowing a comprehensive assessment of their manufacturing precision. The reconstructed models showed good layer-to-layer adhesion, indicating successful integration of individual printing layers to form coherent structures. The imaging also revealed the presence of micropores within the printed splints, likely resulting from the printing process, which could potentially improve the scaffold's biological properties by promoting cell infiltration and nutrient diffusion. Quantitative analysis of the porosity showed that the M1 splint had approximately 71% porosity, while the M2 splint had a slightly higher porosity of approximately 81%. This difference in porosity levels may be due to variations in printing parameters or design specifications between the two splints. In summary, micro- CT imaging showed the structural integrity and porosity of 3D printed Pec-g-PCL splints. The observed micropores and porosity levels provide valuable insights into the scaffold's architecture, which is crucial for optimizing its performance in tissue engineering applications, particularly in tracheal regeneration. As explained above, the 3D polymer splint of the present invention comprising such structural features (micropores and the porosity level) facilitates specific interactions between integrins and ligands, which may affect cell proliferation and further promote optimal cell growth, migration, attachment, and proliferation when employed in-vivo. Thus, the splint of the present invention demonstrates its potential application in tissue engineering. Further, mechanical characterization and Finite Element Analysis (FEA) was carried out to obtain valuable insights into the performance of C-shaped tracheal splints (M1 and M2) of Example 5 when subjected to radial and lateral compression, the results are demonstrated in Figure 7 and 8. Mechanical testing:

[0080] The tracheal splints (M1 and M2) of Example 7 underwent testing using a Mecmism adhesion tester (MultiTest dv (u)). A load cell of 500 N and a strain rate of 1 mm / min were applied during the tests. Mechanical properties were assessed under radial and lateral compression as depicted in Figures 7a to 7f and was compared to commercially available PCL-based splints. Splints measuring 12 mm x 10 mm x 10 mm (length x width x height) were compressed at a constant strain rate of 1 mm / min between flat steel plates. Testing ceased at a strain of 50% of the printed structures. Each specimen underwent at least 6 replicates for compression tests. The trachea demonstrates distinct mechanical characteristics in radial and lateral directions, primarily due to its elliptical lumen shape and the presence of the membranous portion. These structural features significantly affect the trachea's biomechanical behavior and functionality in vivo. Through subjecting the splints to controlled loading conditions, the behavior and response of the said splints, provided an understanding of their mechanical properties and structural integrity. During radial compression tests, it was observed that the load-deformation curves of Pec-g-PCL were consistently lower than those of PCL for both models (Figure 7a-c). This suggests that Pec-g-PCL has reduced stiffness compared to PCL. However, no significant difference in the load- deformation curves or stiffness was observed during lateral compression tests (Figures 7d to f). This suggests that the biodegradable hybrid polymer splint of the present invention could effectively maintain the tracheal shape and withstand physiological pressures in vivo.

[0081] Additionally, to evaluate recovery capability, cyclic compression was conducted under a 30% strain using the same load cell. For a specific strain level, splints underwent 5 cycles of compression, and stress-strain curves were recorded, as depicted in Figure 7f. Splint M2 showed lower stiffness compared to M1 The resilience after deformation is another important feature of external airway splints for medical applications. A radial compression and lateral compression test of 50 cycles was conducted on the tracheal substitutes. The load-displacement evolution curve of M1 and M2 under radial compression is shown in Figures 7g and 7h. After the initial compression cycle, there was a rapid decrease in load. The energy dissipated was higher for M1 compared to M2 under radial compression. Interestingly, the mechanical response of the biodegradable polymer splint of the present invention was maintained at a relatively stable level even when the samples were further compressed under a radial deformation of 60% for 50 cycles.

[0082] Finite Element Analysis (FEA):

[0083] Finite element analysis was further conducted on the tracheal splints to understand their stress concentration under compression (radial and lateral compression) with an identical displacement were calculated via linear elastic simulation using ANSYS workbench. The commercial PCL polymer material properties were assumed isotropic with a Young’s modulus of 124 MPa and a Poisson’s ratio of 0.3 according to previous literature.C3D8 element was selected for meshing with a meshing size of 0.25 mm. The boundary conditions are shown in Figure 8. In radial compression tests, the symmetry plane of the splint was constrained, which can only move vertically. A displacement of 1 mm was applied on the top rigid plate, and the bottom rigid plate was fully fixed. In lateral compression test, the symmetry plane of the splint was fully fixed, and a horizontal displacement of 1 mm was applied to the rigid plate on the right side. A STL mesh is imported into ANSYS Fluent 14.5 (ANSYS Inc., Pennsylvania, USA), where tetrahedron elements and a patch-independent algorithm are utilized for mesh generation. The base mesh size is set to 0.2 mm, but the final size is determined through a mesh-independent evaluation, ensuring a tolerance of less than 0.3%. Depending on the model size, the mesh typically comprises 400,000 to 1,400,000 elements. Inlet velocity of 5m / s, resulting in Reynolds Numbers (Re) between 1542 and 1886. Since Re is below 2000, a laminar model is chosen for simulations. Despite its potential lack of realism in this context, the calculations were conducted using the k - Є turbulence model. The outlet is maintained at a constant pressure of 1.0 atm. The airway wall is assumed to exhibit a noslip boundary condition. Air is treated as a Newtonian fluid with a constant density of 1.225 kg / m3and viscosity of 1.8 x 10-5kg / m-s, which are reasonable assumptions considering the low pressure within the airway.

[0084] As observed from Fig. 8 the stress was mainly concentrated on the top and lateral side under radial compression deformation for both type of porous splints (M1 and M2). However, the maximum stress was found to locate at the corners of the convoluted grooves in case of M1 and corners of hexagonal array of M2 splints. The larger value and distribution area of the maximum stress of M1 splints might be the reason for their earlier yield under radial compression as observed in Figure 7. A similar trend was also observed for M1 and M2 splints under lateral compression. Figure 8 thus suggests the suitability of the splint of the present invention to achieve desired mechanical properties and durability for airway support applications.

[0085] Degradation studies

[0086] Assessing the biodegradability of scaffolds is pivotal for gauging their compatibility with the temporal demands of cell differentiation and tissue regeneration. Following this, the impact of degradation on the characteristics of Pec-g-PCL based scaffold cylindrical scaffold of 5 mm x 5 mm x 6 mm (width x diameter x height) was evaluated by analysing its weight loss, surface morphology, compressive modulus by both radial and lateral compression tests, in comparison to those of commercial PCL scaffolds, (used as controls).

[0087] Figure 9 represents the degradation studies of pectin-grafted polycaprolactone (pectin-g-PCL) scaffolds, under both accelerated (Figures 9a to 9c) and physiological degradation conditions (Figure 9d to 9f). To study the degradation pattern of pectin-g-PCL scaffold of Model 2, a thorough degradation protocol with six replicas of cylindrical scaffolds, measuring 5mm x 5mm x 6mm (width x diameter x height) was prepared. These replicas were then immersed in lOmL of 5M NaOH solution at 37°C for 35 days for accelerated degradation studies. Subsequently, physiological degradation was carried out in phosphate-buffered saline PBS (pH 7.4) under shaking condition at 30°C for up to 6 months. At specific time intervals, the samples were removed, dried overnight at 30°C to remove any remaining liquid, and then weighed to determine weight loss. Additionally, Figure 9b-c and 9e-f depicts the compressive modulus of the scaffold under accelerated and physiological conditions respectively, confirming the correlation between mass loss and mechanical property reduction. The morphological characteristics of the degraded scaffolds were then examined using SEM as depicted in Figure 10. In addition, degradation was assessed using thermal and mechanical evaluation methods.

[0088] Figure 9a, demonstrates a comparison of weight loss between Pec-g-PCL scaffolds of Model 2 and commercial PCL scaffolds. Initially, the degradation process showed non-uniform mass loss, mainly due to cleavage in the non-uniform regions of the polymer. However, over time, the Pec- g-PCL based scaffolds degraded more quickly, with a mass loss of 35±0.4% (Pec-g-PCL) by day 28, compared to the commercial PCL scaffolds, which had a mass loss of 21±0.9%. By day 35, the Pec-g-PCL scaffolds completely broke down, resulting in a loss of mechanical strength. The faster degradation of pectin-g-PCL scaffolds under hydrolytic conditions is caused by several factors. Firstly, the presence of hydrophilic pectin in the PCL scaffold leads to increased water absorption, which accelerates the breakdown of polymer chains. Additionally, the inclusion of glycolide (g) moieties in the graft copolymer makes it more susceptible to hydrolytic breakdown compared to pure PCL scaffolds. This increased breakdown is due to the easily broken ester linkages present in both the pectin and PCL segments of the graft copolymer. As a result, the combination of increased water absorption and ester linkages leads to faster degradation of pectin- g-PCL scaffolds, exceeding the breakdown rate of conventional PCL scaffolds.

[0089] Figures 9b-9c shows the compressive modulus of degraded scaffolds changing over time, confirming the degradation process observed in weight loss studies (Figure 9a). Initially, the mechanical properties of the biodegradable hybrid polymer splint of the present invention remained stable for the first three days. However, after 28 days, there was a noticeable reduction in the elastic modulus for all samples. The Pec-g-PCL scaffolds experienced a 50% decrease, while PCL scaffolds experienced a 40% decrease. Unfortunately, the mechanical properties were unable to be analysed on Day 35 due to scaffold disintegration. This decrease in modulus is likely due to the significant mass loss caused by the accelerated degradation mechanism.

[0090] Figure 9d to 9f shows the scaffolds degradation under physiological conditions representing the human body's pH and temperature (pH 7.4 at 37°C). This environment is important to study because it mimics the conditions in the body where tissue regeneration occurs. Interestingly, as observed from Figures 9d to 9f, even after six months of incubation, no noticeable changes were observed in key parameters such as weight, mechanical strength, or the overall structure of the scaffold. These findings highlight the exceptional stability of the Pec-g-PCL scaffold under physiological conditions. This stability is highly desirable in tissue engineering, as it ensures that the scaffold can provide the necessary support and framework for tissue regeneration over a long period. Therefore, these results suggest that the biodegradable hybrid polymer splint of the present invention could be a reliable structure for supporting tissue regeneration processes in the body.

[0091] Figure 10 shows the SEM analysis used to study the change in the 3D-printed scaffolds as they degraded. The surface morphology of degraded samples was investigated using FEI Nova NanoSEM. The splint were cross-sectioned, mounted on SEM stubs, and sputter-coated with Iridium for analysis. For cell culture samples, a series of steps were followed: after being washed thrice with sterile PBS and once with sodium cacodylate buffer, they were fixed with 100 pL of 1% osmium tetroxide for 30 minutes in the dark. Subsequently, samples were washed with Milli- Q water and then sequentially dehydrated with 50%, 70%, 90%, 95%, and 100% ethanol (10 minutes for each step) before being air-dried overnight. The said Figure confirms that surface erosion is the main way these structures break down under alkaline conditions. Initially, the surface of the scaffold was smooth, but over time it became rougher and developed pores and cracks. This change was more noticeable in Pec-g-PCL scaffolds compared to commercial PCL scaffolds, suggesting that the lower crystallinity and higher water accessibility in Pec-g-PCL scaffolds are the main reasons for their breakdown. Further, the enhanced porosity and surface roughness of degraded pectin-g-PCL scaffolds are beneficial for cell infiltration and tissue integration in biomedical applications. These degradation studies from Figures 9 and 10 clearly show that the breakdown of the fabricated scaffolds is influenced by both their chemical composition and design.

[0092] Biological studies

[0093] In tracheal tissue engineering, the adherence of cells to the scaffold is crucial.

[0094] In vitro cytotoxicity tests and proliferation studies were conducted using the L929 mouse fibroblast cell line to evaluate scaffold potential, which is suggested by ISO 10993 as the preferred cell line for preliminary cytotoxicity testing of biomaterials. L929 cells were grown in DMEM (Gibco) with 10% FBS (Gibco). Polymeric scaffolds were sterilized by UV light for 3 h, soaked and washed with ethanol and ultrapure distilled water. The L929 cells (50,000 cells / ml) were seeded in 24 well plates for 24 hours before the incubation with scaffold of Example 5. Further for cell viability studies, the polymeric scaffolds were incubated with cells in a 5% CO2incubator at 37 °C for 24 hours. To assess the cytotoxicity of the hybrid polymer splint of the present invention, an MTT assay was conducted. For cell proliferation studies, scaffolds seeded with cells were incubated for 1, 3, 5, and 7 days and a resazurin assay was conducted. After incubation, 100 μl of resazurin solution was added followed by further incubation for 4 hours. The solution's optical density (OD) was measured at 550 nm using an absorbance plate reader to assess the cell viability of each sample. The obtained values were compared with the values of plate control.

[0095] Cellular function and interaction of the scaffolds were evaluated by staining Actin filaments by Alexafluor 488 phalloidin and counterstaining cell nuclei with NucBlueTM Live. Briefly L929 cells at a density of 5 x 104were seeded to scaffolds and incubated for 1, 3 & 7 days under standard conditions. On completion of the incubation period scaffold containing cells was washed thrice with PBS fixed with 4% paraformaldehyde solution. Cells were further permeabilized and treated with 5% BSA for 20 minutes. Further cells were stained with 1: 100 dilution of Alexafluor 488 phalloidin (30 min). The epifluorescence images were acquired using an Axio Observer ZICarl Zeiss microscope at excitation / emission wavelengths of 488 / 520 and 358 / 463 nm, respectively.

[0096] As observed from Figure 1 la, the cell viability was determined to be greater than 70%, indicating the non-cy to toxic nature of the developed polymers according to ISO standards.

[0097] Figure 11b and Figure 11c showcases the promising progression of viable cells on the tracheal splints of different geometries from day 1 to day 7. As compared to commercially available PCL splint, the Pec-g-PCL tracheal splint (Model 2) of the present invention demonstrated notably superior cell proliferation compared to others. However, proliferation of cells on commercially available PCL-based splints was significantly hindered by the 7thday, as observed from Figure 1 lb, suggesting that the hydrophobic nature and inadequate pore volume and interconnectivity of the scaffold could impede further proliferation once cell density reaches a certain threshold. Further the SEM analysis (Figure 1 Id to 1 If) of the scaffolds from Figure 11 further supports these observations, showing that Pec-g-PCL splints of the present invention with a more favourable 3D microenvironment facilitated attachment and proliferation. The combined results of the resazurin assay (cell viability and proliferation tests) and SEM images indicate that both surface properties and scaffold architecture of the biodegradable polymer splint of the present invention are crucial factors in enhancing cell viability and proliferation.

[0098] ADVANTAGES OF THE INVENTION

[0099] 1. The biodegradable hybrid polymer splint of the present invention promotes better cellular interaction, supporting the attachment, survival, migration, proliferation, and differentiation of chondrocytes essential for tracheal cartilage development.

[0100] 2. The tracheal splints provided by the present invention are biodegradable and bioresorbable, eliminating the need for additional surgical procedures to remove the splint, thereby reducing patient risk and improving recovery times. 3. The application of 3D printing technology allows for the creation of patient-specific implants with intricate geometries tailored to individual anatomical requirements, enhancing the compatibility and effectiveness of the splint.

[0101] 4. The splint design by 3D printing including precise pore sizes and porosity, matches the mechanical properties of tracheal cartilage, ensuring structural integrity and functionality similar to the natural trachea.

[0102] 5. The splint's design and material composition support the principles of tissue engineering, promoting the natural regeneration and repair of tracheal tissues.

Claims

We Claim:

1. A process for preparation of a biodegradable hybrid polymer splint, wherein the process comprising: a) carrying polymerization reaction of polycaprolactone PCL with pectin in presence of a catalyst to obtain pectin-g-PCL; and b) 3D printing of pectin-g-PCL of step (a) to obtain the biodegradable hybrid polymer splint. wherein the ratio of pectin to polycaprolactone is in a range from 1 :5 to 1 : 1000.

2. The process as claimed in claim 1 , wherein the catalyst is selected from tin-based catalyst ranging from 0.1 mol% to 1 mol%.

3. The process as claimed in claim 1, wherein the polymerization reaction is carried out at a temperature in a range from 100°C to 200°C.

4. The process as claimed in claim 1 , wherein the polymerization reaction is carried out for a period ranging from 16 hours to 48 hours.

5. The process as claimed in claim 1, wherein the 3D printing is carried out by hot melt extrusion.

6. The process of claim 1, wherein the 3D printing is carried out at a temperature in a range from 70°C to 100°C.

7. A biodegradable hybrid polymer splint prepared by process as claimed in claim 1 , wherein said biodegradable hybrid polymer comprising honeycomb structure, rectangular convolutions or combination thereof.

8. The biodegradable hybrid polymer splint as claimed in claim 7, wherein an inner diameter of splint ranging from 5 mm to 15 mm, a length ranging from 10 mm to 20 mm, and a wall thickness ranging from 1 mm to 3 mm.

9. The biodegradable hybrid polymer splint as claimed in claim 7, wherein the porosity of splint is at least 70%.

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