1. Introduction

Figure 1. A) Architecture of blood vessel. Endothelial cells positioned in a single layer on subendothelial extracellular matrix and internal elastic lamina make up the tunica intima. The external elastic lamina, extracellular matrix, and smooth muscle cells make up the tunica media. Fibroblasts and stem/progenitor cells are distributed throughout the connective tissue that makes up the tunica adventitia. B) Procedure for percutaneous coronary intervention. It employs a catheter (a thin flexible tube) to place a stent to open-up blood vessels. The balloon is initially inflated, followed by the expansion of the stent, which pushes the plaque against the artery wall. After the stent has been successfully implanted, the balloon is deflated and removed, leaving the stent to keep open the artery.

Figure 1. A) Architecture of blood vessel. Endothelial cells positioned in a single layer on subendothelial extracellular matrix and internal elastic lamina make up the tunica intima. The external elastic lamina, extracellular matrix, and smooth muscle cells make up the tunica media. Fibroblasts and stem/progenitor cells are distributed throughout the connective tissue that makes up the tunica adventitia. B) Procedure for percutaneous coronary intervention. It employs a catheter (a thin flexible tube) to place a stent to open-up blood vessels. The balloon is initially inflated, followed by the expansion of the stent, which pushes the plaque against the artery wall. After the stent has been successfully implanted, the balloon is deflated and removed, leaving the stent to keep open the artery.

2. Clinical Perspective of Stents

3. Limitations

4. Fabrication Technologies

4.1. 3D Printing

The process of 3D printing involves the creation of three-dimensional things through the sequential deposition of polymeric material, based on a predetermined design generated using computer-aided design software. This technology is primarily utilized for quick prototyping purposes. It is an automated process which requires less time and is thus more efficient than conventional injection molding [86]. Therefore, it is widely used for manufacturing of industrial and medical products [87]. This method offers many advantages for medical applications, including customized production, high precision, and fast fabrication [88]. Several techniques such as fused deposition modelling (FDM), selective laser sintering (SLS), and stereolithography (SLA) have been adopted for 3D printing [89,90].

4.1.1. Extrusion Based 3D Printing

4.1.1.1. Fused Deposition Modelling

Among all the 3D printing techniques, fused deposition modelling (FDM) is the most cost-effective and readily used [91,92]. Materials are initially manufactured into filaments and fed into a heating nozzle for filament-based printing. Later, this filament is melted, extruded, and placed onto a platform to create a three-dimensional structure layer-by-layer [93]. Therefore, the utilization of raw materials is utmost, and this fabrication process does not require any pre-requisite mold or template [94,95].

Misra et al. [96] have utilized this technique to print a flat stent with the traditional polymer PCL and incorporated graphene nanoplatelets to generate a nanocomposite. The overview of their work has been summarized in Figure 2A. A fibrin targeting probe was firstly devised to locate clot in clogged arteries through computed tomography (CT) and accordingly, a computer aided design (CAD) model for the stent to be printed (40 mm in length and 4 mm wide) was prepared. The printed stent exhibited adequate flexibility as shown in Figure 2B(i) and later folded over a mandrel (diameter 2 mm) to achieve the cylindrical structure. Dynamic mechanical analysis represented in Figure 2B(ii) shows a 22% improvement in mechanical strength, with Young’s modulus increasing from 726 ± 50 kPa in PCL stents in comparison to 889 ± 76 kPa for graphene-incorporated PCL stents. These results corroborated with existing literature where graphene has been reported to increase mechanical properties of materials [47]. The efficiency of the clot targeting probe was revealed in cross-sectional CT images where clot treated with targeted probe showed high contrast in Figure 2C(i and ii), in comparison to control solution which produced negligible contrast in Figure 2C(iii and iv). The structure of the graphene containing PCL stent is shown in Figure 2C(v). Finally, the authors had performed an ex vivo feasibility experiment in swine model because of high resemblance of pig with human heart and similar physical dimensions of coronary arteries [97]. After sacrificing a six-month-old swine, the interior dimensions of the coronary artery were determined using a CT scan of the pig's heart, prior to deploying the stent. The 3D-printed stent was implanted at the desired location, followed by CT scanning to corroborate the expansion of the artery as shown in Figure 2C(vi to viii). In the stenting process, the artery was first inflated with a stent loaded on a catheter, and then the stent was moved into the artery with the help of the catheter, as shown step-wise in Figure 2D(i to iv).

Polyesters such as poly(lactic acid) (PLA) and poly(glycolic acid) (PGA) are acceptable polymers for fabrication of coronary stents. In a study by Jia et al. [98], PLA has been used for fabrication of self-expandable vascular stents to make use of the shape memory properties of the polymer [99]. Shape memory polymers can be used to fabricate stents because these could be self-deployed at body temperature without the need of any additional setup. Taking advantage of this property, 3D printing of the vascular stents was performed with a fused deposition modeling (FDM) based printer with nozzle diameter of 0.4 mm. The shape recovery of the stent was done at 70 °C in which it could revert to its initial shape within 5 s. The shape memory property of PLA has also been explored by Wu et al. for fabrication of stents by FDM with Negative Poisson’s Ratio (NPR), as shown in Figure 2E(i to iii) [100]. The effects of geometric parameters on wall thickness, radial compressive property of the stent, and stent diameter were studied. The changes in dimensions of diameter and length of the stent were calculated before and after the shape memory experiment. When the deformation and recovery temperature were both kept at 65 ºC, then the diameter recovery reached to around 95% and length above 97%. The scientists discovered that increasing surface coverage and wall thickness and decreasing stent diameter increased implant radial force per unit length. Thus, this kind of material has great future for self-expandable cardiovascular stent applications. However, stents generally undergo longitudinal foreshortening which means that the implant contracts in terms of its length when dilated. Hence, clinicians must use a stent longer in dimensions than the area to be stented in the artery where the plaque formation has occurred [101]. To circumvent this issue, Wang et al. [102] have developed a novel screw extrusion-based 3D printing system that includes a mini-screw extruder and stents were fabricated with Zero Poisson’s Ratio (ZPR). The effects of surface morphology on the geometric and fabrication parameters were checked for the 3D printed PCL stents. The cell viability of human umbilical vein cells on the stents was 90 ± 5%, which means that these implants can be utilized in vascular applications.

After promising results with PLA and PCL individually as polymers for BVS fabrication, another group used both materials to 3D print stents in a customized 3D printer. Hence, Guerra et al. had utilized a novel tubular 3D printer for rapid manufacture of BVS which was based on FDM [103,104]. The machine methodology has been depicted in Figure 2Fi while 3D printed PCL and PLA stents are shown in Figure 2Fii. Their previous results suggest that this technology could be utilized for BVS manufacturing as the group had managed to manufacture scaffolds under 5 min with up to achieve up to 85% precision [103]. A composite material comprising of PLA and PCL was prepared in a layered fashion with similar molecular weights of the polymers and the resulting blend improved the limitations with the individual material alone. PCL stents presented excellent expansion behavior but high recoil ratios, on the other hand, PLA stents presented excellent recoil ratio with inadequate radial expansion, due to their rigid nature. However, when composite stents were fabricated either with PLA or PCL as the inner layer, the implants possessed the properties of PCL stents (radial expansion) and PLA stents (recoil ratio). This approach could provide a good solution for bioresorbable scaffolds. The dynamical mechanical analysis show that these polymers cannot be used alone with moduli either high or low, whereas the composite groups have values of modulus in the optimum range, making them ideal. Most importantly, the layer arrangement did not affect any of the mechanical property parameters. The composite PCL/PLA stents degraded almost moderately for all layer configurations, mostly owing to PLA degradation. The faster rate at which PLA breaks down would leave a frame made only of PCL in the end. PCL on the outside may be helpful because it may increase the growth of vascular cells. On the other hand, the internal PLA layer may slow the growth of cells, which may help control restenosis. This study was reported to the first work where the group had developed and presented composite PCL/PLA stents using a 3D-printing process based on FDM that could comply with the strict BVS requirements [104].

A similar version of a rotary 3D printing methodology (Figure 3Ai) was utilized by Qui et al. [105] for the fabrication of PCL stents (Figure 3Aii) which were later coated by 2-N, 6-O-sulfated chitosan (26SCS). Their 3D printing machine was based on electrospinning, where under voltage of 4 kV, the PCL filaments were deposited onto a rotatory mandrel. The sulphated (S-PCL) and non-sulphated (PCL) stents were checked for morphology through scanning electron microscopy where PCL stents exhibited a smooth surface in Figure 3C(i to iii) and S-PCL stents showed a rough and porous surface in Figure 3C(iv to vi). This surface topography will ultimately lead to better endothelial cell proliferation [106] and play a major role in blood and cell compatibility. The degradation kinetics of the stents have been depicted in Figure 3D, where weight loss of S-PCL was observed to be 16 and 7% in presence and absence of lysozyme at 60 days, respectively. This result shows that PCL stents can be modified with appropriate non-toxic enzymes to accelerate its slow degradation rate, which may otherwise pose a limitation for using this polymer. A material is considered to be non-hemolytic if rate of hemolysis is less than 2% and acceptable if rate is between 2-5% [107]. Figure 3E shows that hemolysis rate for all samples was below 5%, which suggested the hemocompatible property of both PCL and S-PCL stents. In this study, the in vivo loadings experienced by stents were simulated by performing lateral crush resistance tests (Figure 3F). The mechanical properties were not altered after the 26SCS modification since no significant difference between PCL and S-PCL stents were observed in the force–displacement and stress–strain curves (Figure 3G).

Figure 2. Ai) The presented image depicts a cross-section of an artery situated on the surface of the heart and plaque accumulation inside lumen; ii) A fibrin-targeted iodinated CT contrast probe is used to locate the blood clot and meanwhile, volumetric CT imaging accurately measures the obstruction; iii) Imaging data is used to design a personalized stent using CAD software; iv) PCL-GR polymer composite is printed using an FDM; v) prototyped stent is inserted inside the artery; vi) two drugs are added for sequential release; vii) CT imaging monitors healing process and (vii, ix) Biodegradation of polymer occurs within a wider artery. Bi) Flexible stent (4 cm × 0.4 cm in dimensions) printed using customized extrusion setup and ii) Mechanical properties of 3D printed PCL and PCL-GR stents expressed as stress-strain curve. C(i, ii) Cross-sectional CT images showing clot treated with targeted probe, scale bar – 0.75m and (iii, iv) Cross-sectional CT images showing clot treated with non-targeted probe, scale bar – 0.75m. D(i to iv) Description of the steps involved in implanting a 3D-printed PCL-GR stent in an ex vivo pig artery model. Reproduced with permission from [96] and E) Shape recovery of PLA stents, i) original stent, ii) crimped stent, iii) recovered stent. Reproduced from [100] and Fi) machine methodology and ii) 3D printed PCL (white) and PLA stents (black). Reproduced from [104].

Figure 2. Ai) The presented image depicts a cross-section of an artery situated on the surface of the heart and plaque accumulation inside lumen; ii) A fibrin-targeted iodinated CT contrast probe is used to locate the blood clot and meanwhile, volumetric CT imaging accurately measures the obstruction; iii) Imaging data is used to design a personalized stent using CAD software; iv) PCL-GR polymer composite is printed using an FDM; v) prototyped stent is inserted inside the artery; vi) two drugs are added for sequential release; vii) CT imaging monitors healing process and (vii, ix) Biodegradation of polymer occurs within a wider artery. Bi) Flexible stent (4 cm × 0.4 cm in dimensions) printed using customized extrusion setup and ii) Mechanical properties of 3D printed PCL and PCL-GR stents expressed as stress-strain curve. C(i, ii) Cross-sectional CT images showing clot treated with targeted probe, scale bar – 0.75m and (iii, iv) Cross-sectional CT images showing clot treated with non-targeted probe, scale bar – 0.75m. D(i to iv) Description of the steps involved in implanting a 3D-printed PCL-GR stent in an ex vivo pig artery model. Reproduced with permission from [96] and E) Shape recovery of PLA stents, i) original stent, ii) crimped stent, iii) recovered stent. Reproduced from [100] and Fi) machine methodology and ii) 3D printed PCL (white) and PLA stents (black). Reproduced from [104].

The degradation rate of polymer is an important deciding factor for stent fabrication. Poly(p-dioxanone) (PPDO) is an aliphatic semi-crystalline polyester with ester and ether bonds with a resorption time of 6 months [108]. The material is biocompatible, bioresorbable, and flexible. The Food and Drug Administration has also authorized numerous PPDO medical implants, including suture, clip, staple, pin, and mesh [109]. A recent study explored at how the temperature of the nozzle, the temperature of the bed, the thickness of the layers, and the speed of printing affect the mechanical features of PPDO stents [110]. The enhancement of Young's modulus and yield strength of PPDO FDM parts was observed with an increase in nozzle temperature and bed temperature. The same polymer has been used by another group to fabricate stents using FDM using poly(dioxanone) (PDO) and PLA blends, and the authors implemented a virtual testing system to examine the structural response of custom 3D printing bioresorbable stent designs [111]. Virtual testing is crucial for studying stent mechanics in an in-silico scenario, avoiding costly animal studies and trial-and-error methods.

Figure 3. A) Machine methodology of 3D printing. B) 3D printed PCL stents. C) Scanning electron micrographs of (i to iii) PCL stent and (iv to vi) S-PCL stent with different magnifications. D) Degradation kinetics of S-PCL stents in presence and absence of lysozyme. E) RBC hemolysis percentage after being exposed to PCL and S-PCL stent extracts at different time points. F) Testing apparatus and inset showing diagram for lateral crush resistance testing. G) Stress-strain curve of PCL and S-PCL stents. Reproduced from [105]. H) ABBOTT BVS1.1, PALMAZ-SCHATZ, and ART18Z, flat-shaped CAD models with finally printed stents, scale bar – 5 mm. I) Scanning electron micrographs depicting the adhesion of human blood platelets onto surface of stents printed using different designs. Diagram of the force-displacement relationship for (J) radial compression test and (K) three-point bending test. L) Stent recoil expressed as percentage and M) stent diameter while the balloon expands. Reproduced with permission from [112].

Figure 3. A) Machine methodology of 3D printing. B) 3D printed PCL stents. C) Scanning electron micrographs of (i to iii) PCL stent and (iv to vi) S-PCL stent with different magnifications. D) Degradation kinetics of S-PCL stents in presence and absence of lysozyme. E) RBC hemolysis percentage after being exposed to PCL and S-PCL stent extracts at different time points. F) Testing apparatus and inset showing diagram for lateral crush resistance testing. G) Stress-strain curve of PCL and S-PCL stents. Reproduced from [105]. H) ABBOTT BVS1.1, PALMAZ-SCHATZ, and ART18Z, flat-shaped CAD models with finally printed stents, scale bar – 5 mm. I) Scanning electron micrographs depicting the adhesion of human blood platelets onto surface of stents printed using different designs. Diagram of the force-displacement relationship for (J) radial compression test and (K) three-point bending test. L) Stent recoil expressed as percentage and M) stent diameter while the balloon expands. Reproduced with permission from [112].

4.1.1.2. Solvent Cast 3D Printing

For melt-based techniques, processing pressures and temperatures need to be maintained at high value so that continuous flow is maintained. This presents a limitation because the type of polymer that can be utilized for printing becomes restricted. A second type of extrusion-based printing is solvent-cast 3D printing (SC-3DP), which offers an alternative approach to print a wider range of polymers. In this case, polymers are first dissolved in a volatile solvent and while a solid polymer filament gets deposited, the solvent gets evaporated. However, since stents have many interconnected mesh structures, printing on a flat substrate by applying this technique becomes challenging. Therefore, a cylindrical substrate could overcome these difficulties while developing SC-3DP. This technique was utilized by Singh et al. [112] where they had utilized polycaprolactone stents reinforced with carbonyl iron powder (CIP) according to previously existing designs such as ABBOTT BVS 1.1, PALMA-SCHATZ and ART18Z. The CAD designs and the printed stents have been shown in Figure 3H. The scanning electron micrographs show that the stents possess excellent hemocompatibility as seen by inactivated adhesion of platelets in all the three types, although the least adherence was noticed in PALMAZ-SCHATZ stent (Figure 3I). The ART18Z stents and ABBOTT BVS 1.1 displayed a higher (6.59 ± 0.52 N) and lower (3.32 ± 0.36 N) value of radial compression force respectively, with moderate value (4.65 ± 0.21 N) for PALMAZ-SCHATZ stent (Figure 3J). The flexibility in case of PALMAZ-SCHATZ stent was also moderate compared to ART18Z stents and ABBOTT BVS1.1 (Figure 3K). This radial compressive force is inversely related to the stent’s recoil ratio, since stents with highest radial compression force will experience least recoil, as shown in Figure 3L. These results are also corroborating with the stent diameter calculated after over-expansion, since ART18Z expanded the least on over-inflating with balloon (Figure 3M). The enhancement of this hemocompatibility and mechanical properties could be attributed to the addition of 1 wt% of CIP in PCL matrix [113]. Overall, the results concluded that the stents printed with PALMAZ-Schatz design had better biological and mechanical properties than ABBOTT BVS1.1 and ART18Z that could be used in the treatment of stenotic arteries. An advancement to this work was carried out by the same group where the effect of process parameters on flexibility of stents, radial compression load, and percentage reduction in strut width and thickness, was evaluated [114]. The printing speed and layer thickness were observed to be important factors and contributed to the flexibility, radial compression load and reduction in strut width and thickness of the composite stents. The interaction of layer thickness and printing speed also influenced the mechanical properties of the printed stents. It was observed that at slow printing speed and small layer thickness, the maximum load for bending and radial compression load were obtained. The optimized process parameters were 1.08% CIP concentration, layer thickness around 0.2 mm and printing speed 8.02 mm/s. By following the parameters, the output responses were observed to be 1.13 N load for bending, 4.12 N radial compression load, 12.80% reduction in strut width, and 1.90% reduction in strut thickness. For future work, these optimized parameters would be beneficial in design and fabrication of stents with varied topologies.

4.1.2. Light-Based Printing

While various research groups were working on the traditional polymers, Lith et al. [115] had thought of developing a technology for on-spot fabrication of high-resolution bioresorbable vascular stents based on a versatile form of additive manufacturing technology that will be entirely based on patient-specific design. In their endeavors, a four step process will be followed for placement of stents in patients: 1) vascular parameters of the patient to be assessed with imaging techniques, 2) the design of the stent to be made with computer-aided software based on patient-specific parameters, 3) the stent is fabricated on-the-spot by 3D printing following the design, and 4) the customized stent can now be utilized for implantation in patient. The authors had utilized an antioxidant, photocurable, and bioresorbable citrate-based biomaterial and printing was carried out using a custom-made micro-continuous liquid interface production system (mCLIP), a technology already described by Tumbleston et al. [116] It is based on the principle that oxygen can either quench the photoexcited photo initiator or create peroxides by combining with the free radical from the photo initiator, creating a “dead zone” that cannot be photocured. If stereolithography is conducted above an oxygen-permeable window, a continuous liquid interface will be produced by a thin liquid layer, which will be due to oxygen contact between the window and the cured part. Previously, Park’s group had used bio-plotting that does not accommodate stent customization [117], whereas Flege’s BVS exhibited resolution and poor surface finish [118]. Most importantly, both techniques require fabrication times of 10 h or more. On the other hand, this new mCLIP technology was a continuous process that could speed up the printing process tremendously as compared to FDM that works on layer-by-layer assembly. Hence, this group introduced the first successful mCLIP-produced, antioxidant and customizable BVS that is similar to the mechanical strength of nitinol stent. The group used methacrylated polydiolcitrates which were photo-cured, and intrinsic antioxidant properties were also utilized to chelate trace transition metals that could be potentially harmful. This property was lacking in the conventional polymers used in 3D printing. This work was further optimized by Ware et al. [119] in which they could fabricate a 2 cm tall vascular stent comprising of 4000 layers in 26.5 min. In another study by Oliveira et al. [120], the same polymer has been utilized for printing BVS with another light-assisted 3D printing technique, known as digital light processing (DLP). Commercial low cost DLP printers also provide good resolution, which are comparable to expensive 3D printers that employ mCLIP and projection micro-stereolithography techniques. Nitric oxide (NO) is released by endothelial cells which inhibits adhesion of platelets, maintains vascular tone, and inhibits SMC proliferation [121]. The researchers employed the nitric oxide (NO) donor known as S-nitroso-N-acetyl-d-penicillamine (SNAP) and integrated it into the methacrylated poly(dodecanediol citrate) (mPDC) scaffolds via absorption. The SEM images show the ultrastructure of the stent (3 mm in diameter) with increasing magnification (Figure 4A(i to iii)), where Figure 4A(i) shows that the stent has a consistent geometric structure characterized by clearly defined boundaries and a textured surface with indentations. The layers exhibit strong adhesion and maintain their structural integrity without any surface imperfections, as depicted in Figure 4A(ii and iii). Since polyesters hydrolyze slowly under physiological conditions, taking over 3 years, the mPDC stents were subjected to accelerated degradation (temperature 60 °C and pH 12) and percentage mass loss and change in pH have been shown for 21 days. The mass loss reached 20% after 5 days of rapid degradation, remained stable until day 10, then increased to 60% at day 15, and remained consistent throughout the experiment. This mass loss pattern indicates a two-step hydrolytic breakdown process. During mPDC hydrolysis, the pH fluctuation reduces practically linearly towards zero due to the partial neutralization of the basic solution (pH 12) by free citric acid. Further, after approximately 2 h, different variants of mPDC/SNAP scaffolds attain a constant NO release rate and this value falls within the range of NO release from healthy endothelial cells (Figure 4B). The mechanical properties of the stents were further evaluated, where Figure 4C presents the compressive stress-strain curves derived from the compression testing of 3D printed stents with a diameter of 6 mm. These stents were tested in their uncured state and subsequently subjected to post-cure periods of 3, 6, and 12 min. The viscoelastic behavior of stents is evident when subjected to linear compressions up to 50% of their original diameter, a property commonly observed in elastomeric materials. The optimal post-cure irradiation time for the current mPDC stents was determined to be 3 min. This duration resulted in consistent stress-strain behavior and an elastic response of up to 50% strain. This elasticity enables the stent structure to recover after deployment. On the other hand, Figure 4D illustrates the stress-strain behavior of stents with diameters of 4 mm, 5 mm, and 6 mm that underwent a post-cure period of 3 min. In a similar manner, it was observed that stents with diameters ranging from 5 to 6 mm, subjected to a post cure duration of 3 min, exhibited superior compressive capabilities compared to stents with a diameter of 4 mm, for the intended use. It can also be seen that the stent could transform into a fully collapsed state on compression and after removal of the mechanical load, it immediately reverts back to its original expanded state, as represented in Figure 4E(i to iii). Thus, these stents are an advancement in material design for fabrication of bioresorbable vascular stents, where the NO release will play an important role as a therapeutic entity. However, there were some shortcomings of these studies, such as the materials used for fabricating the stents are not FDA approved and hence require extensive studies before they can be taken further.

Figure 4. A(i to iii) Scanning electron micrographs of 3D printed with increase in magnification from left to right, demonstrating the printing accuracy from the millimeter to the micrometer length scale. B) Real time NO released by 3D-printed mPDC/80, 40 and 20 discs, after being immersed in PBS (pH 7.4) and 37 oC. Stress-strain compression curves for C) uncured and post-cured mPDC stents at 3, 6, and 12 min and D) mPDC stents with 4-, 5-, and 6-mm diameters post-cured for 3 min and E) Images of a 6 mm stent i) before, ii) fully compressed, and iii) extended after compression. Reproduced with permission from [120].

Figure 4. A(i to iii) Scanning electron micrographs of 3D printed with increase in magnification from left to right, demonstrating the printing accuracy from the millimeter to the micrometer length scale. B) Real time NO released by 3D-printed mPDC/80, 40 and 20 discs, after being immersed in PBS (pH 7.4) and 37 oC. Stress-strain compression curves for C) uncured and post-cured mPDC stents at 3, 6, and 12 min and D) mPDC stents with 4-, 5-, and 6-mm diameters post-cured for 3 min and E) Images of a 6 mm stent i) before, ii) fully compressed, and iii) extended after compression. Reproduced with permission from [120].

5. Functionalization with Bioactive Molecules

Figure 5. Different approaches for surface modification of stents to enhance endothelialization.

Figure 5. Different approaches for surface modification of stents to enhance endothelialization.

6. Conclusions

7. Future Perspectives

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