2.2.2. Characterisation of Polymer Blends

#### Fourier Transform Infrared (FTIR)

FTIR spectra were obtained in a reflective absorbance mode on a Perkin Elmer spectrometer (Waltham, MA, USA) with a constant spectral resolution of 4 cm−<sup>1</sup> , in the range of 4000 to 550 cm−<sup>1</sup> , and after 16 scans. The reported spectra were analysed quantitatively using Perkin Elmer Software (Waltham, MA, USA) version 10.

#### Differential Scanning Calorimetry (DSC)

Thermal property analyses were carried out using a Differential Scanning Calorimetry (Mettler Toledo, Greifensee, Switzerland). In standard aluminium pans, about 10 mg of the samples were from room temperature to 250 ◦C at a rate of 20 ◦C/min and held isothermally for 5 min to exclude all previous thermal history. After cooling at 5 ◦C/min to −80 ◦C at a rate of 10 ◦C/min, samples were heated again at 20 ◦C/min to 250 ◦C. All of the experiments were conducted under a nitrogen atmosphere.

#### Mechanical Testing

The mechanical properties of all blends were conducted in tensile uniaxial mode using an Instron universal testing machine model 3366 (Norwood, MA, USA) at a crosshead speed of 5 mm/min according to ASTM D638. The samples analysed were 60 mm × 5 mm in size, with a thickness of around 1.0 mm. The slope of the straight-line stress-strain curve was used to calculate the Young's Modulus (YM) of the polymer blends and the effects of tensile strength and percentage of elongation at break. The mean and standard deviation of five measures were used to calculate all results.

#### Scanning Electron Microscope (SEM)

SEM images were captured using a tabletop SEM (Hitachi, Tokyo, Japan). The crosssectional surfaces were obtained from a tensile examination of a broken dumbbell. Before microscopy experiment, all specimens were sputtered with a thin layer of gold.

#### Water Absorption Study

The water absorption test was performed on 10 mm square samples according to ASTM D570. The samples were dried at 60 ◦C for 24 h to achieve a stable weight. The dried samples were weighed and placed in the test plate wells, prewetted with PBS solution before filling with 5 mL of PBS, pH 7.4 at 37 ◦C. The samples were incubated and tested at 8, 24, 48 and 72 h. The excess water was carefully removed with tissue paper, and the samples were re-weighed. Water absorption was calculated based on the amount of water absorbed according to Equation (1):

$$\text{Water absorption (\%)} = \frac{\text{Wwt} - \text{Wdry}}{\text{Wdry}} \times 100\tag{1}$$

For each composition, three specimens were examined to achieve an average value.

#### In Vitro Degradation Study

PBS soaking tests were used to mimic the hydrolytic degradation activity of TPU/PLA blends and control scaffolds. The samples were weighed after drying overnight at 60 ◦C. Each sample was individually enclosed in a plastic container filled with a 1X PBS solution and incubated at 37 ◦C in an orbital shaker (Stuart S1500, Illinois City, IL, USA) at a shaking rate of 50 rpm. Every 7 days, PBS was refreshed, and the test lasted up to 7 months. The samples were rinsed three times with purified water before being dried overnight and weighed at each time point. Equation (2) was used to measure the weight loss of the materials.

$$Depradation\left(\%\right) = \left(1 - \frac{W\_n}{W\_o}\right) \times 100\tag{2}$$

*W*<sup>0</sup> is the initial sample weight, and *W<sup>n</sup>* is the weight of the same sample after degradation for a time, *n*.

#### In Vitro Biocompatibility Assay

The blended samples were evaluated for biocompatibility and possible use as tissue engineering scaffolds in biomedical applications.

#### Pellet Sterilisation

Prior to in vitro testing, the TPU/PLA pellets were sterilised by being immersed in 70% ethanol (*v/v*) for 2 h, followed by rinsing three times with 1x PBS to eliminate all traces of ethanol. The pellets were then air-dried in a sterile atmosphere before being sterilised for 2 h with ultraviolet light. This phase ensured that any pollutants on the surface of the pellets were removed.

#### Pesto Blue Viability Assay

In order to be considered for biomedical applications, any material must not impose any toxicity to the surrounding tissues. The toxicity of the TPU and PLA was investigated in this study and was defined as a reduction in cell growth to less than 50% viability. The sterile pellets were immersed in complete α-MEM overnight before testing. BEAS-2-B (human bronchial epithelial cells) were purchased form American Type Culture Collection (ATCC, Manassas, VA, USA) were grown to confluence in complete α-MEM containing 10% FBS and 1% antibiotic–antimycotic (AA) solution at 37 ◦C in an incubator with 5% carbon dioxide (CO2). In 24-well plates, one pellet was placed in each well, followed by direct cell seeding on top of the pellet at a seeding number of 1 <sup>×</sup> <sup>10</sup><sup>4</sup> cells per well, and the plates were incubated in a CO<sup>2</sup> incubator for 3 days. As a positive control, a tissue culture plate with fresh α-MEM was used. The toxicity test was performed using Presto Blue Cell Viability Reagent on days 1, 2 and 3. The metabolic product of viable cells was released into the culture medium, reduced resazurin to resorufin, and changed the colour from blue to pinkish red. When reduction did not occur in a nonviable setting, the blue-coloured resazurin was preserved. An automatic ELISA reader was used to test the fluorescence of the samples in triplicate. Equation (3) was used to measure the cell toxicity percentage:

$$\text{Cell trajectory (\%)} = \frac{\text{Fluorescence\\_treatment} - \text{Fluorescence\\_control}}{\text{Fluorescence\\_control}} \times 100\tag{3}$$

#### Statistical Analysis

The mean and standard deviation of all the results were calculated (SD). The oneway analysis of variance (ANOVA) was used for normally distributed data, while for non-normally distributed data, the nonparametric test was used. The Tukey's test was then used to assess the data's particular differences, with *p* < 0.05 indicating statistical significance. The data were analysed using GraphPad Software (Prism 9.0, GraphPad Software, La Jolla, CA, USA).

#### **3. Results and Discussion**

Apart from being biocompatible and biodegradable, as the foundation of the tracheal structure, the material of the scaffold should have adequate mechanical strength and flexibility to enable physiological function during breathing. The rigid components of the cartilage retain the trachea lumen open and prevent its collapse under negative air pressure, preventing airflow limitations [49]. The purpose of this study is to determine the physical and mechanical properties, as well as the absorption, in vitro degradation, and biocompatibility, of a series of TPU and PLA blends that will be produced as filaments and 3D printed as a potential material for tracheal replacement.

#### *3.1. Fabrication of Filament and Identification of the Materials 3.1. Fabrication of Filament and Identification of the Materials*

printed as a potential material for tracheal replacement.

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The physical blending of the material through the melt extrusion technique of TPU and PLA was chosen in this study to produce filaments that will be fabricated as tracheal scaffold using a 3D printing technique. Both materials were mixed based on their weight percentage and characterised accordingly. FTIR transmission spectra of the samples are presented in Figure 1 with main characteristic bands of pure TPU appeared at 3328 cm−<sup>1</sup> , 2935 cm−<sup>1</sup> , and 2850 cm−<sup>1</sup> which correspond to stretching of -NH- in urethane and asymmetric and symmetric vibrations in -CH2- respectively [50–53]. In addition, 1700, 1531 and 1314 cm−<sup>1</sup> bands were associated with bending and stretching of amide I, II and III bonds. The intensity of characteristic bands in TPU reduces as the PLA contents increases and vice versa. The characteristic bands of PLA were seen at 1750 cm−<sup>1</sup> , 1456 cm−<sup>1</sup> and 1180 cm−1, corresponding to asymmetric vibration of -C=O and the asymmetric and symmetric stretching of -C-O-C bonds, respectively. The characteristic bands and their activities are summarised in Table 1. Overall, the spectra revealed no new chemical bonds, suggesting that both PLA and TPU were successfully compounded during melt blending [54,55]. Physical blending does not alter the chemical characteristic of the constituents of the polymers [34]. Hence, the properties of the blend can be conveniently customized using different compositions of the polymer to suit its application. The physical blending of the material through the melt extrusion technique of TPU and PLA was chosen in this study to produce filaments that will be fabricated as tracheal scaffold using a 3D printing technique. Both materials were mixed based on their weight percentage and characterised accordingly. FTIR transmission spectra of the samples are presented in Figure 1 with main characteristic bands of pure TPU appeared at 3328 cm−1, 2935 cm−1, and 2850 cm−1 which correspond to stretching of -NH- in urethane and asymmetric and symmetric vibrations in -CH2- respectively [50–53]. In addition, 1700, 1531 and 1314 cm−1 bands were associated with bending and stretching amide I, II, and III bonds. The intensity of characteristic bands in TPU reduces as the PLA contents increases and vice versa. The characteristic bands of PLA were seen at 1750 cm−1, 1456 cm−1, and 1180 cm−1, corresponding to asymmetric vibration of -C=O and the asymmetric and symmetric stretching of -C-O-C bonds, respectively. The characteristic bands and their activities are summarised in Table 1. Overall, the spectra revealed no new chemical bonds, suggesting that both PLA and TPU were successfully compounded during melt blending [54,55]. Physical blending does not alter the chemical characteristic of the constituents of the polymers [34]. Hence, the properties of the blend can be conveniently customized using different compositions of the polymer to suit its application.

preventing airflow limitations [49]. The purpose of this study is to determine the physical and mechanical properties, as well as the absorption, in vitro degradation, and biocompatibility, of a series of TPU and PLA blends that will be produced as filaments and 3D

**Figure 1.** FTIR spectra of TPU, PLA, and TPU/PLA blends. **Figure 1.** FTIR spectra of TPU, PLA, and TPU/PLA blends.

**Table 1.** Characteristic bands and corresponding activities of TPU and PLA. **Table 1.** Characteristic bands and corresponding activities of TPU and PLA.


1085 symmetric stretching of -C-O-C- PLA

#### *3.2. Morphology and Miscibility of the Blends 3.2. Morphology and Miscibility of the Blends*

The SEM images of the fractured surface TPU/PLA dumbbell films were analysed to study the phase morphology of the samples. In contrast, the DSC findings were used to determine the thermal properties and miscibility of the samples. As shown in Figure 2, the fractured surface morphology of pure TPU and pure PLA reveals the homogeneous distribution of fibrous TPU and continuous matrix of PLA, respectively. In contrast, in blended compositions, the smooth-edged PLA domains are distributed in the fibrous TPU matrix, resulting in a notably heterogeneous two-phase structure. Both 90/10 and 80/20 samples displayed more fibrous TPU than 70/30 and 60/40 samples, reflecting the flexibility of the blends. However, several separated phase domains with minute debonded holes were also observed, indicating a weak interfacial contact between the TPU matrix and the PLA domain. As the PLA ratio increased, larger and more PLA spheres were detected in the TPU matrix, although they were dispersed equally. All blended polymers exhibit phase separation, suggesting that TPU and PLA are immiscible. The SEM images of the fractured surface TPU/PLA dumbbell films were analysed to study the phase morphology of the samples. In contrast, the DSC findings were used to determine the thermal properties and miscibility of the samples. As shown in Figure 2, the fractured surface morphology of pure TPU and pure PLA reveals the homogeneous distribution of fibrous TPU and continuous matrix of PLA, respectively. In contrast, in blended compositions, the smooth-edged PLA domains are distributed in the fibrous TPU matrix, resulting in a notably heterogeneous two-phase structure. Both 90/10 and 80/20 samples displayed more fibrous TPU than 70/30 and 60/40 samples, reflecting the flexibility of the blends. However, several separated phase domains with minute debonded holes were also observed, indicating a weak interfacial contact between the TPU matrix and the PLA domain. As the PLA ratio increased, larger and more PLA spheres were detected in the TPU matrix, although they were dispersed equally. All blended polymers exhibit phase separation, suggesting that TPU and PLA are immiscible.

**Figure 2.** Cross-section SEM images of reactively extruded films: (**a**) 100/0, (**b**) 0/100. Homogeneous matrices are present in pure TPU and pure PLA (**c**) 90/10, (**d**) 80/20, (**e**) 70/30, (**f**) 60/40. Some fibrous TPU is present in the polymer blends, with PLA domains are dispersed in TPU matrices in all blends. Red arrows show PLA particles in the TPU matrix, while blue arrows indicate the fibrous TPU of the fractured surface. Scale bar of 100 μm. **Figure 2.** Cross-section SEM images of reactively extruded films: (**a**) 100/0, (**b**) 0/100. Homogeneous matrices are present in pure TPU and pure PLA (**c**) 90/10, (**d**) 80/20, (**e**) 70/30, (**f**) 60/40. Some fibrous TPU is present in the polymer blends,with PLA domains are dispersed in TPU matrices in all blends. Red arrows show PLA particles in the TPU matrix, while blue arrows indicate the fibrous TPU of the fractured surface. Scale bar of 100 µm.

TPU is composed of diisocyanate hard segments and polyester macroglycol soft segments, while PLA is an aliphatic polyester. Due to their structures, two glass transition temperature (Tg) and melting temperature (Tm) values were observed in pure TPU, which correspond to hard and soft segments as shown in DSC curves in Figure 3. On the other hand, only one Tg and Tm were seen in pure PLA. TPU is composed of diisocyanate hard segments and polyester macroglycol soft segments, while PLA is an aliphatic polyester. Due to their structures, two glass transition temperature (Tg) and melting temperature (Tm) values were observed in pure TPU, which correspond to hard and soft segments as shown in DSC curves in Figure 3. On the other hand, only one T<sup>g</sup> and T<sup>m</sup> were seen in pure PLA.

**Figure 3.** DSC secondary melting curves of pure TPU and PLA, and TPU/PLA blends. The blue arrows show the first and second Tg values of pure TPU, while red arrows indicate both Tm values of pure TPU. Grey arrows show the Tg and Tm **Figure 3.** DSC secondary melting curves of pure TPU and PLA, and TPU/PLA blends. The blue arrows show the first and second T<sup>g</sup> values of pure TPU, while red arrows indicate both T<sup>m</sup> values of pure TPU. Grey arrows show the T<sup>g</sup> and T<sup>m</sup> values of pure PLA. Finally, the green arrows depict all three Tg values of TPU/PLA blends.

values of pure PLA. Finally, the green arrows depict all three Tg values of TPU/PLA blends. The immiscibility of the blends is further demonstrated by DSC curves and summarized data as shown in Table 2. The Tg values can determine the extent of blend miscibility, partial miscibility, or total immiscibility as a function of the polymer blend composition. Normal Tg values of the TPU and PLA ranges from around −18 °C to −47 °C and 61 °C to 67 °C, respectively [56,57]. Miscible polymers typically have a single Tg, whereas immiscible blends transition temperatures shift toward each other to a degree depending on the mutual miscibility of the phases [58–61]. In the event of a completely immiscible blend system, the blend components may remain their original Tg values, regardless of the blend composition [34]. In this study, the appearance of three distinct Tg values in 90/10, 80/20, 70/30, and 60/40 samples suggest that TPU and PLA are immiscible. Two Tg values correspond to the TPU structure, whereas another Tg value comes from the PLA. The immiscibility was demonstrated by considerable macroscopic phase separation in the blends in SEM. Moreover, all mixed samples that confirm the immiscibility also show two different melting temperatures (Tm). This result is in line with [51,61–63]. In immiscible blends, the properties of the component polymers integrate in such a way that the blend morphology The immiscibility of the blends is further demonstrated by DSC curves and summarized data as shown in Table 2. The T<sup>g</sup> values can determine the extent of blend miscibility, partial miscibility, or total immiscibility as a function of the polymer blend composition. Normal T<sup>g</sup> values of the TPU and PLA ranges from around −18 ◦C to −47 ◦C and 61 ◦C to 67 ◦C, respectively [56,57]. Miscible polymers typically have a single Tg, whereas immiscible blends transition temperatures shift toward each other to a degree depending on the mutual miscibility of the phases [58–61]. In the event of a completely immiscible blend system, the blend components may remain their original T<sup>g</sup> values, regardless of the blend composition [34]. In this study, the appearance of three distinct T<sup>g</sup> values in 90/10, 80/20, 70/30, and 60/40 samples suggest that TPU and PLA are immiscible. Two T<sup>g</sup> values correspond to the TPU structure, whereas another T<sup>g</sup> value comes from the PLA. The immiscibility was demonstrated by considerable macroscopic phase separation in the blends in SEM. Moreover, all mixed samples that confirm the immiscibility also show two different melting temperatures (Tm). This result is in line with [51,61–63]. In immiscible blends, the properties of the component polymers integrate in such a way that the blend morphology is a direct representation of the component morphology. As a result, blend morphology is regarded as a good indicator of blend miscibility.


is a direct representation of the component morphology. As a result, blend morphology is

regarded as a good indicator of blend miscibility. **Table 2.** Differential scanning calorimetry data for TPU, PLA, and blends.

Tg = glass transition temperature, Tm = melting temperature.

T<sup>g</sup> = glass transition temperature, T<sup>m</sup> = melting temperature.

## *3.3. Mechanical Properties*

*3.3. Mechanical Properties*  It is well recognised that the morphology of the materials greatly influences the mechanical properties of blends, and the properties of the final product can be achieved by It is well recognised that the morphology of the materials greatly influences the mechanical properties of blends, and the properties of the final product can be achieved by adjusting the morphology. The tensile strength and YM measurement were derived from the stress–strain curve of the tensile test ASTM D638. Figure 4 depicts the mechanical

properties of the TPU/PLA blends. Pure PLA has the highest tensile strength and YM at 46.48 ± 5.51 MPa and 2282.80 ± 95.60 MPa, preceded by TPU blends with higher PLA contents. This means that the higher the PLA concentration, the stiffer the blend. On the other hand, pure PLA demonstrated the least amount of flexibility because it cannot elongate by more than 5% due to its inelastic nature. In comparison, pure TPU showed the highest flexibility (up to more than 25%) without fracture, as shown in Figure 4b. The flexibility decreases with decreasing TPU concentration, with the 60/40 blend exhibiting the least flexibility of all blends. The morphology of TPU and PLA blends appeared to be related to their mechanical properties, with stretched fibrous TPU can be seen in fractured surfaces images of the 90/10 and 80/20 samples reflecting the blends' flexibility. properties of the TPU/PLA blends. Pure PLA has the highest tensile strength and YM at 46.48 ± 5.51 MPa and 2282.80 ± 95.60 MPa, preceded by TPU blends with higher PLA contents. This means that the higher the PLA concentration, the stiffer the blend. On the other hand, pure PLA demonstrated the least amount of flexibility because it cannot elongate by more than 5% due to its inelastic nature. In comparison, pure TPU showed the highest flexibility (up to more than 25%) without fracture, as shown in Figure 4b. The flexibility decreases with decreasing TPU concentration, with the 60/40 blend exhibiting the least flexibility of all blends. The morphology of TPU and PLA blends appeared to be related to their mechanical properties, with stretched fibrous TPU can be seen in fractured surfaces images of the 90/10 and 80/20 samples reflecting the blends' flexibility. properties of the TPU/PLA blends. Pure PLA has the highest tensile strength and YM at 46.48 ± 5.51 MPa and 2282.80 ± 95.60 MPa, preceded by TPU blends with higher PLA contents. This means that the higher the PLA concentration, the stiffer the blend. On the other hand, pure PLA demonstrated the least amount of flexibility because it cannot elongate by more than 5% due to its inelastic nature. In comparison, pure TPU showed the highest flexibility (up to more than 25%) without fracture, as shown in Figure 4b. The flexibility decreases with decreasing TPU concentration, with the 60/40 blend exhibiting the least flexibility of all blends. The morphology of TPU and PLA blends appeared to be related to their mechanical properties, with stretched fibrous TPU can be seen in fractured surfaces images of the 90/10 and 80/20 samples reflecting the blends' flexibility.

adjusting the morphology. The tensile strength and YM measurement were derived from the stress–strain curve of the tensile test ASTM D638. Figure 4 depicts the mechanical

adjusting the morphology. The tensile strength and YM measurement were derived from the stress–strain curve of the tensile test ASTM D638. Figure 4 depicts the mechanical

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**Figure 4.** Mechanical properties of TPU/PLA blends showing (**a**) tensile strength and Young's modulus in MPA, and (**b**) percentage of elongation at break. **Figure 4.** Mechanical properties of TPU/PLA blends showing (**a**) tensile strength and Young's modulus in MPA, and (**b**) percentage of elongation at break. **Figure 4.** Mechanical properties of TPU/PLA blends showing (**a**) tensile strength and Young's modulus in MPA, and (**b**) percentage of elongation at break.

As a substitute for the cartilage wall, the scaffold is expected to possess adequate strength to keep the airway open and resist collapse. At the same time, it maintains the flexibility to allow flexion/bending despite intrathoracic pressure differences during breathing cycles [63,64]. TPU is known to have flexibility and is widely used in soft tissue engineering [23,24,65,66]. The present study revealed that the tensile strength and Youngs' modulus of TPU were proportionately increased when PLA was added. In contrast, the percent elongation reduces accordingly. The results obtained were similar to those reported by Lis–Bartos et al. (2018) [67] and Mi et al., (2013) [51]. The overall physicomechanical behaviour of immiscible systems is critically dependent on two demanding structural parameters. First, an appropriate interfacial tension leads to a phase size that is small enough to allow the material to be considered macroscopically homogeneous. Secondly, an interphase adhesion is strong enough to assimilate stresses and strains without disrupting the established morphology [68]. Even though the blend was not miscible, it is noted that they considered having good compatibility with each other. This is due to the composition of diisocyanate hard segments and polyester macroglycol soft segments of the TPU and aliphatic polyester in PLA, which forms hydrogen bonding between the molecules of the blend, as shown in Figure 5. As a substitute for the cartilage wall, the scaffold is expected to possess adequate strength to keep the airway open and resist collapse. At the same time, it maintains the flexibility to allow flexion/bending despite intrathoracic pressure differences during breathing cycles [63,64]. TPU is known to have flexibility and is widely used in soft tissue engineering [23,24,65,66]. The present study revealed that the tensile strength and Youngs' modulus of TPU were proportionately increased when PLA was added. In contrast, the percent elongation reduces accordingly. The results obtained were similar to those reported by Lis–Bartos et al. (2018) [67] and Mi et al., (2013) [51]. The overall physicomechanical behaviour of immiscible systems is critically dependent on two demanding structural parameters. First, an appropriate interfacial tension leads to a phase size that is small enough to allow the material to be considered macroscopically homogeneous. Secondly, an interphase adhesion is strong enough to assimilate stresses and strains without disrupting the established morphology [68]. Even though the blend was not miscible, it is noted that they considered having good compatibility with each other. This is due to the composition of diisocyanate hard segments and polyester macroglycol soft segments of the TPU and aliphatic polyester in PLA, which forms hydrogen bonding between the molecules of the blend, as shown in Figure 5. As a substitute for the cartilage wall, the scaffold is expected to possess adequate strength to keep the airway open and resist collapse. At the same time, it maintains the flexibility to allow flexion/bending despite intrathoracic pressure differences during breathing cycles [63,64]. TPU is known to have flexibility and is widely used in soft tissue engineering [23,24,65,66]. The present study revealed that the tensile strength and Youngs' modulus of TPU were proportionately increased when PLA was added. In contrast, the percent elongation reduces accordingly. The results obtained were similar to those reported by Lis–Bartos et al. (2018) [67] and Mi et al., (2013) [51]. The overall physicomechanical behaviour of immiscible systems is critically dependent on two demanding structural parameters. First, an appropriate interfacial tension leads to a phase size that is small enough to allow the material to be considered macroscopically homogeneous. Secondly, an interphase adhesion is strong enough to assimilate stresses and strains without disrupting the established morphology [68]. Even though the blend was not miscible, it is noted that they considered having good compatibility with each other. This is due to the composition of diisocyanate hard segments and polyester macroglycol soft segments of the TPU and aliphatic polyester in PLA, which forms hydrogen bonding between the molecules of the blend, as shown in Figure 5.

**Figure 5.** Schematic diagram of hydrogen bonding between the molecules of PLA and TPU. Reproduced with permission from Feng and Ye, Journal of Applied Polymer Science; published by John Wiley and Sons, 2011. [69]. **Figure 5.** Schematic diagram of hydrogen bonding between the molecules of PLA and TPU. Reproduced with permission from Feng and Ye, Journal of Applied Polymer Science; published by John Wiley and Sons, 2011. [69]. **Figure 5.** Schematic diagram of hydrogen bonding between the molecules of PLA and TPU. Reproduced with permission from Feng and Ye, Journal of Applied Polymer Science; reprinted with permission from ref [69], Copyright 2011 John Wiley and Sons.

The blend is suitable for the fabrication of a tissue-engineered scaffold since the mechanical properties are within the normal human tracheal range: tensile stiffness 1–15 MPa [70] and YM 12.2–20.5 MPa in young and old humans; ~16 MPa in both human male and female [64].

The purpose of combining two polymers is to produce a material with improved physical properties over the parent polymers. The degree of modification is relative to the quantity of each composition of the polymer. The morphology of the blend and the dispersion of the phases have a considerable effect on the thermomechanical characteristics. For example, TPU with a lower tensile strength is combined with PLA to increase the tensile strength. Despite the modest reduction in elongation at break, the final material can be adjusted by varying the polymer composition to attain the optimal combination of thermal and mechanical properties for specific applications.

#### *3.4. Water Absorption and In Vitro Degradation Rate*

An ideal implant material should possess optimal absorption and biodegradable properties similar to the regenerative process of the native tissue [71]. This fills up surrounding tissues and provides a biocompatible framework that allows cells, blood vessels, and newly created tissues to develop and maintain an extracellular matrix [72].

The swelling behaviour of a material (or its ability to absorb water) is an important factor to consider when constructing a scaffold. Although excessive water absorption damages the scaffold's morphology, its absence causes inadequate absorption, inhibiting cell growth in vivo [54]. The absorption rate was determined by immersing the samples in PBS solution. Pure PLA showed the highest absorption rate in PBS solution, which was approximately 5% in the first hour, decreased significantly after 8 h, and remained stable until 48 h. On the other hand, other compositions exhibited nearly equal absorption, varying from 1 to 2 percent, but decreased steadily and remained nearly stable until 48 h. No significant differences were found in any of the compositions. The material's absorption property indicates the scaffold's ability to bind to the surrounding metabolites and promotes the transportation of nutrients and cell integration throughout the scaffold [73].

One of the primary goals of biodegradable tissue-engineered scaffolds is to provide a mechanical structure or framework to support the extracellular matrix (ECM). At the same time, it degrades evenly/slowly via hydrolytic degradation, allowing the surrounding tissue to recover the supporting function of the scaffold [74]. The monomeric components of the polymers are innocuously eliminated through the body system [75]. In the degradation study over 7 months, all of the samples showed a gradual loss with pure TPU (100/0) exhibiting the fastest degradation rate, followed by 90/10, 80/20, 70/30, 60/40 and pure PLA (0/100). The higher the TPU concentration, the faster rate of degradation. The degradation percentage up to 7 months was less than 6%, suggesting that the blended material is long-term stable. In contrast, the degradation of PLA is slower, which started only after 1 month of incubation and gradually degraded at a very slow rate. Up until 7 months, the degradation percentage was approximately 2% and was the lowest among other blends. Figure 6 shows both the rate of absorption and degradation results.

**Figure 6.** (**a**) Water absorption rate of TPU, PLA, and the blends, (**b**) In vitro degradation study of the samples. Pure TPU **Figure 6.** (**a**) Water absorption rate of TPU, PLA, and the blends, (**b**) In vitro degradation study of the samples. Pure TPU shows the fastest degradation rate, while the slowest is that of pure PLA.

shows the fastest degradation rate, while the slowest is that of pure PLA. The biodegradation of polymeric biomaterials mainly depends on hydrolytic degradation of polymer chains. For example, degradation of polyurethane occurs when water molecules infiltrate the polymer network, triggering hydrolysis of the polyurethane chains, including the chemical dissolution of ester and amide bonds [76]. Similarly, the hydrolytic degradation of PLA started by the breaking of the ester link of the polymeric chain [77]. Hydrolytic degradation has been described as belonging to two types: surface erosion and bulk erosion. Surface erosion occurs exclusively at the polymer–water interface, while bulk erosion occurs uniformly throughout the polymer [76,77]. However, it is critical to emphasise that this degradation investigation was conducted on nonporous film specimens without plasma, biomolecules, or cells. The rate of degradation is important, as it is controlled by various other factors in vivo, which significantly enhances the deg-The biodegradation of polymeric biomaterials mainly depends on hydrolytic degradation of polymer chains. For example, degradation of polyurethane occurs when water molecules infiltrate the polymer network, triggering hydrolysis of the polyurethane chains, including the chemical dissolution of ester and amide bonds [76]. Similarly, the hydrolytic degradation of PLA started by the breaking of the ester link of the polymeric chain [77]. Hydrolytic degradation has been described as belonging to two types: surface erosion and bulk erosion. Surface erosion occurs exclusively at the polymer–water interface, while bulk erosion occurs uniformly throughout the polymer [76,77]. However, it is critical to emphasise that this degradation investigation was conducted on nonporous film specimens without plasma, biomolecules, or cells. The rate of degradation is important, as it is controlled by various other factors in vivo, which significantly enhances the degradation of polymers.

radation of polymers. Erosion and bulk degradation are the two possible mechanisms through which polymers degrade. Crystallinity and chain orientation of polymers are particularly essential in degradation processes. The degradation process begins with water diffusion into the amorphous regions, followed by random hydrolytic scission of the ester bonds. After degrading the amorphous sections, the hydrolytic attack continues to the crystalline structures. Due to the loose packing of the molecules in the amorphous area of a polymer, they are more susceptible to attack by reactive species or solvents than the molecules in the crystalline region [78–80]. TPU's amorphous structure is the possible cause that results in a faster degradation rate than semicrystalline PLA, as seen by its lower glass transition Erosion and bulk degradation are the two possible mechanisms through which polymers degrade. Crystallinity and chain orientation of polymers are particularly essential in degradation processes. The degradation process begins with water diffusion into the amorphous regions, followed by random hydrolytic scission of the ester bonds. After degrading the amorphous sections, the hydrolytic attack continues to the crystalline structures. Due to the loose packing of the molecules in the amorphous area of a polymer, they are more susceptible to attack by reactive species or solvents than the molecules in the crystalline region [78–80]. TPU's amorphous structure is the possible cause that results in a faster degradation rate than semicrystalline PLA, as seen by its lower glass transition temperature (Tg).

#### temperature (Tg). *3.5. In Vitro Biocompatibility*

*3.5. In vitro Biocompatibility*  The viability of cell culture is shown by the proportion of viable cells in a population. A decrease in the percentage of viable cells below 50% of total cell growth is described as the toxicity of the material towards the cells. BEAS-2B was cultured for three days on ster-The viability of cell culture is shown by the proportion of viable cells in a population. A decrease in the percentage of viable cells below 50% of total cell growth is described as the toxicity of the material towards the cells. BEAS-2B was cultured for three days on sterilised TPU, PLA, and TPU/PLA scaffolds to test their biocompatibility and toxicity of the material, and the result is shown in Figure 7.

ilised TPU, PLA, and TPU/PLA scaffolds to test their biocompatibility and toxicity of the

material, and the result is shown in Figure 7.

**Figure 7.** Presto Blue viability assay towards BEAS-2B cells. **Figure 7.** Presto Blue viability assay towards BEAS-2B cells.

Both pure TPU and PLA indicated biocompatibility towards BEAS-2B cells. After 24 h of incubation, cell proliferation was noted, exhibiting viability of cells was more than 100% in both pure materials. The result showed that the viability of BEAS-2B was greater than 80% in all compositions up until day 3, indicating that none of the compositions were toxic to the cells. Pure PLA outperformed pure TPU and other blends in terms of viability. TPU has been proven to be biocompatible and suitable for regenerative medicine, either alone or in conjunction with other polymers. Therefore, this result was expected [17,81,82]. This study is also consistent with the findings by Harynska and colleagues (2018) [83]. Similarly, the U.S. Food and Drug Administration (FDA) has authorised PLA as a biodegradable and biocompatible polymer for application in the human body due to its absorbability and nontoxicity [33,35,73,84]. Both pure TPU and PLA indicated biocompatibility towards BEAS-2B cells. After 24 h of incubation, cell proliferation was noted, exhibiting viability of cells was more than 100% in both pure materials. The result showed that the viability of BEAS-2B was greater than 80% in all compositions up until day 3, indicating that none of the compositions were toxic to the cells. Pure PLA outperformed pure TPU and other blends in terms of viability. TPU has been proven to be biocompatible and suitable for regenerative medicine, either alone or in conjunction with other polymers. Therefore, this result was expected [17,81,82]. This study is also consistent with the findings by Harynska and colleagues (2018) [83]. Similarly, the U.S. Food and Drug Administration (FDA) has authorised PLA as a biodegradable and biocompatible polymer for application in the human body due to its absorbability and nontoxicity [33,35,73,84].

#### **4. Conclusions**

**4. Conclusions**  Due to the shortage of organs and tissues for organ transplantation, synthetic materials are some of the treatment alternatives for trachea replacement, which necessitates developing an ideal material with desirable properties. Apart from biocompatibility and nontoxicity in the body system, mechanical properties are some of the important factors that must be considered due to the anatomy of the trachea in the body. TPU and PLA were selected for this study because of their inherent properties and wide use in biomedical applications. However, the properties of these materials limit their application. Polymer blending is an appealing and cost-effective method for developing new material with improved properties by combining physically existing polymers instead of synthesizing entirely new polymeric materials. This study investigated the mechanical properties, water absorption, biodegradability, and biocompatibility of melt blended TPU and PLA polymers as a material of choice in 3D printed tracheal tissue engineering. It was demonstrated that both materials were successfully compounded. Even though all blend compositions were not miscible, their morphology and mechanical properties showed improvement for the proposed use. Furthermore, the blended material is biocompatible and has an appropriate absorption and degradation rate, making it viable for use as a filament material in additive manufacturing for medical purposes. Therefore, the polymer blending of TPU and PLA can be a cost-effective strategy to improve the properties of TPU and PLA and Due to the shortage of organs and tissues for organ transplantation, synthetic materials are some of the treatment alternatives for trachea replacement, which necessitates developing an ideal material with desirable properties. Apart from biocompatibility and nontoxicity in the body system, mechanical properties are some of the important factors that must be considered due to the anatomy of the trachea in the body. TPU and PLA were selected for this study because of their inherent properties and wide use in biomedical applications. However, the properties of these materials limit their application. Polymer blending is an appealing and cost-effective method for developing new material with improved properties by combining physically existing polymers instead of synthesizing entirely new polymeric materials. This study investigated the mechanical properties, water absorption, biodegradability, and biocompatibility of melt blended TPU and PLA polymers as a material of choice in 3D printed tracheal tissue engineering. It was demonstrated that both materials were successfully compounded. Even though all blend compositions werenot miscible, their morphology and mechanical properties showed improvement for the proposed use. Furthermore, the blended material is biocompatible and has an appropriate absorption and degradation rate, making it viable for use as a filament material in additive manufacturing for medical purposes. Therefore, the polymer blending of TPU and PLA can be a cost-effective strategy to improve the properties of TPU and PLA and produce filaments for 3D printing.

produce filaments for 3D printing. **Author Contributions:** Conceptualization, A.A.S. and B.H.Y.; methodology, A.A.S., B.H.Y., Z.A.A.H., and M.J.; validation, A.A.S., B.H.Y. and Z.A.A.H.; formal analysis, A.A.S., B.H.Y. and Z.A.A.H.; investigation, A.A.S.; resources, A.A.S., B.H.Y. and Z.A.A.H.; data curation, A.A.S. and B.H.Y.; writing—original draft preparation, A.A.S.; writing—review and editing, A.A.S., B.H.Y. and Z.A.A.H.; visualization, A.A.S., B.H.Y., Z.A.A.H. and M.J.; supervision, B.H.Y., Z.A.A.H. and M.J.; project administration, A.A.S. and B.H.Y.; funding acquisition, A.A.S. and B.H.Y. All authors have **Author Contributions:** Conceptualization, A.A.S. and B.H.Y.; methodology, A.A.S., B.H.Y., Z.A.A.H. and M.J.; validation, A.A.S., B.H.Y. and Z.A.A.H.; formal analysis, A.A.S., B.H.Y. and Z.A.A.H.; investigation, A.A.S.; resources, A.A.S., B.H.Y. and Z.A.A.H.; data curation, A.A.S. and B.H.Y.; writing—original draft preparation, A.A.S.; writing—review and editing, A.A.S., B.H.Y. and Z.A.A.H.; visualization, A.A.S., B.H.Y., Z.A.A.H. and M.J.; supervision, B.H.Y., Z.A.A.H. and M.J.; project administration, A.A.S. and B.H.Y.; funding acquisition, A.A.S. and B.H.Y. All authors have read and agreed to the published version of the manuscript.

read and agreed to the published version of the manuscript.

**Funding:** This research was funded by the Ministry of Higher Education Malaysia for the Prototype Research Grant Scheme (PRGS) with Project Code PRGS/1/2021/SKK07/USM/02/1 and the Nippon Sheet Glass Foundation for Materials Science and Engineering (304/CIPPT/650932).

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** Not applicable.

**Acknowledgments:** The authors would like to thank the Ministry of Higher Education Malaysia and the International Islamic University Malaysia for their support through the Academic Staff Training Scheme fellowship (ASTS) for Asmak Abdul Samat. The authors would like to thank to staff in Regenerative Medicine Laboratory, Advanced Medical and Dental Institute (IPPT), Universiti Sains Malaysia for their assistant and supports.

**Conflicts of Interest:** All authors declare no conflict of interest. The authors alone are responsible for the content and writing of the article.

## **References**

