Structural Design and Mechanical Properties of Metal Vascular Stents Fabricated via Laser Powder Bed Fusion

Abstract

Against the backdrop of high cardiovascular disease prevalence, personalized vascular stents tailored to complex vascular structures have emerged as a critical research focus in interventional therapy. While laser powder bed fusion (L-PBF) offers a novel approach to customized stent manufacturing, systematic evaluations of its forming accuracy and mechanical performance remain insufficient. This study designed a novel vascular stent featuring double-period unequal-height support rings interconnected by M-shaped struts and monolithically fabricated it from 316L stainless steel using L-PBF. Surface quality was enhanced via electropolishing. Balloon expansion, radial compression, and microtensile testing quantitatively characterized stent expansion performance (radial recoil rate, axial shortening rate), radial strength, and intrinsic material mechanical properties. Results demonstrated that L-PBF achieved unsupported fabrication of fine features (150 μm wall thickness); electropolishing significantly improved surface quality, reducing dimensional deviation from 46.7% to 3.3%. The stent exhibited a radial strength of 840 mN/mm and a remarkably low radial recoil rate of 1.37%. Material yield strength (232 MPa) reached 74.1% of that of wrought 316L alloy, with a fracture elongation of 16.44%, ensuring structural integrity during expansion. However, the axial shortening rate (5.56%) exceeded expectations, primarily due to geometric deviation in high-curvature connecting struts suppressing their compensatory mechanism. The M-type stent demonstrated superior radial support via its asymmetric-height support ring design, confirming L-PBF’s feasibility for producing personalized metal vascular stents. Future work necessitates optimization of connecting strut curvature to control axial shortening and integrated topological design with targeted process optimization to advance clinical translation.

1. Introduction

Driven by evolving lifestyles and accelerating population aging, cardiovascular diseases (CVDs) have emerged as the leading cause of mortality nationwide, with an estimated CVD death rate of 334.77 per 100,000 individuals. Percutaneous coronary intervention (PCI) has become the mainstream treatment for atherosclerotic stenosis due to its minimally invasive nature and clinical efficacy. Within this context, vascular stents serve as pivotal PCI devices that restore blood flow by mechanically supporting stenosed vessel segments, simultaneously inhibiting elastic recoil. Conventional metallic stents typically feature a periodic lattice structure composed of radial-support rings interconnected by strut links, collectively governing device deliverability through tortuous vasculature and postimplantation stability.

Vascular stents currently employed in clinical practice primarily comprise metallic and polymeric materials, each exhibiting significant performance trade-offs. Metallic stents (e.g., magnesium alloys, cobalt–chromium alloys, nitinol) exhibit superior mechanical properties, enabling enhanced radial strength at smaller structural dimensions and greater suitability for sustaining high cyclic vascular loads. − However, studies reveal inherent limitations: magnesium alloy stents, despite favorable biodegradability that mitigates long-term foreign body reactions, suffer from excessively rapid degradation rates (<6 months) leading to premature mechanical failure; the superelasticity of NiTi may induce chronic vascular irritation that potentiates restenosis. Polymeric materials are predominantly used for bioresorbable stents, yet their inferior strength and flexibility necessitate larger strut widths and wall thicknesses compared to metallic counterparts. Consequently, increased wall thickness adversely impacts hemodynamics, promoting platelet adhesion, plasma protein deposition, and in-stent restenosis. While biodegradable polymers like poly­(l-lactic acid) (PLLA) demonstrate excellent biocompatibility, their inadequate radial strength (merely 1/5 to 1/10 of metals) and pro-inflammatory degradation byproducts remain critical concerns. Therefore, material selection requires concerted optimization of mechanical performance, biological responses, and hemodynamic compatibility to maximize clinical outcomes.

Conventional metallic vascular stents are predominantly manufactured via laser micromachining, with alternative methods including braiding and injection molding. However, as a material-reduction manufacturing process, this technique inevitably results in significant material waste. Moreover, its geometric freedom is inherently constrained by the two-dimensional planar patterns cut from a single tubular precursor, posing considerable challenges in fabricating complex three-dimensional structures, such as customized bifurcated stents. In comparison to laser-cut vascular stents, additive manufacturing (AM) technologies offer notable advantages for stent production: , Monolithic forming capability, which saves material and shortens production cycles, eliminates internal stress concentrations, and enhances structural integrity; patient-specific customization facilitates tailored stent dimensions (length, diameter) and configurations based on lesion morphology; exceptional geometric freedom enables fabrication of complex topological architectures. Given these attributes, AM has emerged as the most promising approach for manufacturing personalized metallic vascular stents.

As a core additive manufacturing technology, L-PBF has demonstrated significant potential for metallic vascular stent fabrication. Demir and Previtali successfully fabricated CoCr alloy stents via L-PBF, validating its capability for producing fine-featured complex structures. In 2019, Finazzi et al. further monolithically manufactured bifurcation-specific stents using L-PBF, with balloon expansion tests confirming their structural soundness. Jamshidi’s group systematically investigated the parametric sensitivity of L-PBF on NiTi stent accuracy, while Zhang et al. established an inverse correlation between volumetric energy density and surface roughness in L-PBF-processed stents, demonstrating electropolishing’s efficacy for surface enhancement. Safdel et al. further demonstrated that L-PBF can fabricate patient-specific NiTi stents with complex geometries and excellent mechanical properties, including high density, superior compressional recoverability, and remarkable superelasticity. Despite these advances, systematic evaluations of critical mechanical properties (e.g., radial strength, fatigue resistance) in L-PBF-fabricated stents remain underinvestigated.

However, the widespread application of L-PBF in stent manufacturing is constrained not only by issues of surface quality and dimensional accuracy but also by inherent gaps in the mechanical properties of the printed materials. Compared to conventional wrought materials, L-PBF-fabricated metallic materials often exhibit reduced yield strength and significantly lower fracture elongation. Furthermore, the unique microstructures resulting from rapid solidification may lead to anisotropic mechanical behavior and potential brittleness under complex loading conditions. The insufficient ductility raises concerns about the risk of brittle fracture during stent expansion, while compromised radial strength could undermine the essential function of resisting vascular recoil. Therefore, in addition to optimizing process parameters, innovative structural design becomes crucialnot only to compensate for the mechanical deficiencies of the material but also to redistribute stress more effectively, thereby fully unleashing the potential of additive manufacturing in the customization of personalized stents.

The topological configuration of vascular stents serves as the fundamental carrier for functional performance, directly determining mechanical properties, hemodynamic behavior, and manufacturability. − Within periodic lattice structures, radial-support ringsas primary load-bearing elementsgovern radial strength through geometric parameters (e.g., waveform amplitude). Conversely, strut configurations (e.g., straight or N-shaped) modulate axial shortening and flexibility. Wall thickness (t) must balance mechanical and hemodynamic requirements (optimal range: 100–150 μm); excessive thickness enhances radial support at the cost of aggravated flow disturbances and elevated thrombosis risk. , Concurrently, L-PBF process constraints necessitate avoiding elongated horizontal overhangs to prevent thermally induced warpage from rapid melt pool solidification. These structural elements collectively influence core clinical mechanical performance, evaluated through two critical metrics: Expansion performance (quantified via balloon expansion tests: radial recoil indicates wall apposition; axial shortening affects lesion coverage completeness); radial strength (measured by radial compression tests: force per unit length maintains vessel patency). Complementarily, finite element analysis (FEA) enables multiobjective design optimization by simulating stent–balloon–vessel coupling models. These predict plastic strain distributions (critical threshold >30% for fracture risk) and provide theoretical validation.

Building upon this research foundation, to holistically assess the feasibility of L-PBF for monolithic metallic stent fabrication and resultant mechanical performance, this study designed and manufactured a novel M-type vascular stent via L-PBF. Its core innovations comprise: Double-period unequal-height radial-support rings optimizing radial force distribution; M-shaped struts interconnecting dual-ring structures. Medical-grade 316L stainless steel powder was utilized for L-PBF fabrication. Postprocessing included electropolishing to reduce surface roughness, followed by balloon expansion, radial compression, and dog-bone tensile testing. Systematically evaluating the M-type stent’s mechanical properties through expansion performance and radial strength metrics establishes theoretical foundations for clinical translation of personalized vascular stents.

2. Methodology

2.1. Vascular Stent Structural Design

2.1.1. Support Ring Architecture

The stent in this study employs a hybrid structure combining radial-support rings and strut connectors. Existing research confirms that geometric parameters of both elements significantly influence mechanical performance, particularly radial strength. − Literature analysis further reveals that the number of circumferential support units (N) and their arc length (L) critically govern radial behavior. For instance, Zhang et al. demonstrated an inverse correlation between radial strength and increases in N and L; furthermore, optimal stiffness was achieved at N = 6 (even number). Wei et al. pioneered an unequal-height support structure in polymeric stents, which enhances radial strength by enlarging the interunit angle postexpansion. Additionally, reverse-aligned ring configurations exhibit superior mechanical performance compared to sequential arrangements.

Building upon this foundation, we developed a novel metallic stent architecture integrating reverse-aligned configuration, unequal-height geometry, and rotational symmetry. This design was designated as the M-type stent based on its distinctive morphological signature. The radial-support rings were engineered as follows: Sinusoidal unequal-height waveform with dual-period amplitudes; six circumferential support units (N = 6) ensuring symmetric distribution of major/minor support elements while maximizing radial force; reverse-aligned symmetric arrangement between adjacent rings; uniform wall thickness of 150 μm across all rings. The 2D expanded view and unit cell dimensions are detailed in Figure .

1.

Structure and size of the M-type vascular stent support ring expansion diagram.

2.1.2. Strut Configuration

The stent incorporates M-shaped curved struts alternating with linear struts to interconnect adjacent radial-support rings. While the primary function of struts is ring connectivity, their geometric configuration significantly governs axial shortening and flexibility. To analyze mechanical behavior, a representative support unit with attached struts was selected (Figure a). During balloon expansion, pressure P acts on the stent’s luminal surface, inducing circumferential forces F _t at the support unit ends (Figure b). The force-pressure relationship is defined as

$$F_t=PS/2\sin \frac{\alpha }{2}$$ 1

where

2.

Schematic of force distribution on the stent support unit: (a) Three-dimensional structure of the support unit; (b) force analysis on a single support ring in top view; (c) force analysis on a simplified support unit; (d) force analysis on the attached struts.

F_t = Circumferential force at support unit ends (N)

P = Balloon pressure on the intraluminal surface (MPa)

S = Luminal surface area of the support unit (mm²)

α = Central angle of the support unit arc (°)

To establish the relationship between balloon pressure P and tensile force in struts F_t , the support unit was simplified to a planar isosceles triangular model (Figure c). Here, F _t1 = F_t cos­(α/2), where F _t1 denotes the circumferential force in the simplified model. This force can be resolved into orthogonal components: F ₁: radial expansion force; F ₂: axial shortening force that reduces support unit height. The resultant strut tension F _T is given by

$$F_T=F_2=F_{t1}\sin \beta =PS\sin \beta \cot \frac{\alpha }{2}/2$$ 2

where

F _T = Tensile force transmitted to radial-support rings by struts (N)

β = Apex angle of the support unit (°)

The resultant tensile force acting on the strut is denoted as $F_{\mathrm{T}}^{\prime }$ , which constitutes an action-reaction pair with F _T as illustrated in Figure d. The M-shaped curved strut design plays a critical role: under tensile loading $F_{\mathrm{T}}^{\prime }$ , the curved geometry undergoes elastic elongation, thereby compensating for axial shortening caused by support unit height reduction. This mechanism effectively mitigates overall stent axial contraction. Similarly, when subjected to bending loads, the curved strut configuration reduces constraining reaction forces on radial-support rings, permitting greater relative displacement and consequently enhancing stent flexibility.

The three-dimensional stent configuration is illustrated in Figure a. To ensure uniform circumferential support at both terminals postdeployment and prevent vascular wall injury due to stress concentrations, the proximal and distal radial-support rings feature a constant-height geometry. Additionally, six support feet were specifically designed at the stent base to enhance L-PBF manufacturability. Critical geometric parameters comprise an outer diameter of 2.4 mm, uniform wall thickness of 150 μm, consistent ring/strut width of 150 μm, and total length of 19.37 mm. Given the interdependence between stent dimensions and L-PBF-induced material properties, a customized tensile specimen was engineered to mitigate size-effect deviations (Figure b). This specimen possesses a 5 mm gauge length with cross-sectional dimensions (150 × 150 μm) precisely mimicking the stent ring geometry.

3.

Structural schematics of the stent and tensile specimen:(a) 3D stent; (b) custom tensile specimen.

2.2. L-PBF Fabrication of Metallic Vascular Stents

2.2.1. Powder Material

Medical-grade 316L stainless steel powder (HLPOWDER, Hunan, China) was employed, with chemical composition detailed in Table . Scanning electron microscopy (Hitachi SU8200, Tokyo, Japan) revealed highly spherical morphology with minimal satelliting (Figure ), indicating exceptional flowabilitycritical for uniform powder layering in L-PBF. BT-9300ST laser diffraction analysis (Bettersize Instruments Ltd., Dandong, China) showed a particle size distribution spanning D10 (20.70 μm) to D90 (52.39 μm). The powder exhibited: Apparent density: 4.13 g/cm³; tap density: 4.63 g/cm³; Hausner ratio (HR): HR = tap density/bulk density = 1.12. HR < 1.25 confirms favorable flowability, consistent with SEM observations. The high tap density (>4.5 g/cm³) predicts powder bed relative density >50%, substantially reducing porosity in L-PBF-fabricated stents and mitigating strut cracking risks.

1. Chemical Composition of 316L Stainless Steel Powder (wt %).

Si	Cr	Ni	Mn	Mo	S	P	C	O	Fe
0.56	16.52	10.36	0.93	2.47	0.007	0.011	0.015	0.055	Bal.

4.

Scanning electron microscopy image of 316L stainless steel powder.

2.2.2. L-PBF Processing

The stent 3D model designed in SolidWorks 2022 (Dassault Systèmes, France) was imported into a DiMetal-100E metal additive manufacturing system (LASEADD, Guangzhou, China), where concentric contour scanning strategy with 20 μm layer thickness was implementeddirecting a continuous laser beam (140 W power, 40 μm spot diameter) along predefined paths from inner to outer contours at 900 mm/s scan speed under continuous nitrogen inert atmosphere to suppress oxide formation.

2.2.3. Electropolishing

L-PBF-fabricated stents exhibited inherent surface irregularities including melt pool striations and partially melted particles (initial roughness Ra > 8 μm). To achieve implant-grade surface finish (Ra< 0.5 μm), electrochemical polishing was performed using a P-016 electropolishing system (Pudeng, Taiwan, China) with 316L stents as anodes and high-purity lead plates as cathodes immersed in ECP-316L proprietary electrolyte. A constant DC voltage of 7 V maintained current stability at 2 A (current density: 4 A/cm²) under constant-voltage mode. The process was conducted at 40 ± 2 °C for 90 s with continuous stent rotation at ≈30 rpm.

2.3. Tensile Testing

The radial collapse resistance of vascular stents is governed by the yield strength (YS) and ultimate tensile strength (UTS) of the strut material, while flexibility and fatigue life are predominantly governed by the elastic modulus (E) and elongation at break (δ). To precisely characterize the intrinsic mechanical properties of the strut structures fabricated via L-PBF, uniaxial tensile tests were performed on electrolytically polished specimens using a high-precision micromechanical testing system (ElectroForce 3220-AT, TA Instruments, USA), as illustrated in Figure .

5.

High-precision micromechanical testing system: (a) Overall setup of the system; (b) tensile fixture.

2.4. Balloon Expansion Testing

Balloon expansion testing of the vascular stent was performed using electrolytically polished specimens. The experimental setup incorporated a PB-20E-0630 single-use balloon dilatation catheter and a BI-20A-10 inflation device (both from Changmei Medtech, Changzhou, China). The dilation process was monitored and recorded using an HW4K stereomicroscope (Zhiwei Precision, Shenzhen, China). A schematic of the experimental setup is shown in Figure a. To eliminate external interference and ensure uniform pressure distribution on the stent’s inner surface, the stent was secured using a custom-designed adjustable expansion platform (Figure b). Balloon pressure was incrementally increased from 0 to 8 bar in 0.5 bar steps. After each pressure increment, the stent outer diameter was measured at three locations: proximal, midbody, and distal segments. The stent expansion ratio was calculated as follows:

$$r_x=d(p)/d_0$$ 3

where d(p) is the outer diameter measured at pressure P, and d ₀ is the initial outer diameter of the stent.

6.

(a) Schematic diagram of the balloon expansion testing setup; (b) custom-designed adjustable expansion platform.

The M-type vascular stent was expanded to its nominal diameter of 3.6 mm. The stent dimensions were recorded using the HW4K stereomicroscope in the as-fabricated state, at full expansion (3.6 mm), and after balloon deflation. Subsequently, the radial recoil rate and axial shortening rate of the stent were calculated using eqs and .

$$R_{\mathrm{recoil}}=(1-\frac{D_{\mathrm{deflated}}}{D_{\mathrm{inflated}}})\times 100\%$$ 4

where:

R _recoil radial recoil rate,

D _deflatedouter diameter of the stent measured after balloon deflation,

D _inflatedouter diameter of the stent measured at full balloon expansion (target diameter: 3.6 mm).

$${\gamma }_{\mathrm{axial}}=(1-\frac{L_{\mathrm{deflated}}}{L})\times 100\%$$ 5

where:

γ_axialaxial shortening rate,

L _deflatedstent length measured after balloon deflation,

Lstent length in the as-fabricated (unexpanded) state.

2.5. Radial Compression Testing

Radial compression testing was conducted in accordance with the ISO 25539–2:2014 standard to evaluate the stent’s ability to resist circumferential pressure exerted by the vessel wall. The test was performed at 37 °C using a high-precision micromechanical testing system ElectroForce 3220-AT (TA Instruments, New Castle, USA) equipped with flat plate compression fixtures (Figure ). The stent, after electropolishing and expansion to 3.6 mm, was compressed at a constant rate of 0.15 mm/min until a 50% diameter reduction was achieved. The load–displacement curve recorded in real-time was used to calculate radial strength. As specified by the standard, to ensure comparability of radial support performance across stents of different lengths, the radial force was normalized. The expression for the normalized radial strength is

$$P_{\mathrm{load}}=\frac{F}{L_{\mathrm{deflated}}}$$ 6

where:

7.

Flat plate compression fixture.

P _Loadcompressive load,

Fpressure value of the tested stent,

L _deflatedstent length measured after balloon deflation.

3. Results and Discussion

3.1. Surface Quality

Figure presents the macroscopic morphology of the vascular stent and tensile specimens as-fabricated directly on the substrate via L-PBF. The stent arrays exhibit regular arrangement without collapse or fracture, indicating that unsupported monolithic fabrication was achieved through optimized geometric design, ensuring manufacturing precision.

8.

As-fabricated stents and tensile specimens on the substrate.

Figure compares the surface morphology evolution of the stent before and after electropolishing via SEM analysis. In the as-fabricated state (Figure a, b), the stent lattice structure retained its geometric integrity without cracking. However, the surface exhibited significant roughness characterized by two predominant defect types: adhered sintered particles and scaled melt pool morphology. This indicates that while the initial metallic stent validates the capability of L-PBF for monolithic fabrication, its surface quality was unacceptably low with high roughness. Analysis attributes the sintered particles to incompletely melted fine metal powder particles (typically <50 μm in size) originating from the edges of laser scan paths or within the heat-affected zone; these ″satellite particles″ adhere to the surface due to insufficient thermal input. The scaled melt pool morphology results from the synergistic effects of: (1) rapid solidification shrinkage (cooling rate >10⁶ K/s), (2) melt pool instability induced by Marangoni convection and recoil pressure, and (3) secondary remelting caused by overlapping laser scans. This fluctuating solidification pattern is directly replicated onto the component surface. ,,− Concurrently, the filamentary nature of the stent structures (diameter <200 μm) constrains melt pool size and heat conduction, accentuating edge effects and promoting balling and stair-stepping phenomena. Furthermore, regions such as support contact points, steep downward-facing surfaces, and complex overhangs exhibited pronounced roughness due to compromised heat dissipation and powder adhesion. Elevated surface roughness on medical implants poses dual hazards: ,, it promotes abnormal platelet adhesion, compromising biocompatibility, and serves as a fatigue crack initiation site, reducing service life. Consequently, these surfaces fail to meet the ASTM F2477 surface finish requirement (Ra < 0.5 μm) for medical implants. To enhance surface quality, electrochemical polishing (ECP) was applied to the fabricated stents. Figure c, d demonstrates that this process selectively removes surface-sintered particles (removal rate >95%) while preserving lattice geometry stability, achieving an optimal balance between functional surface enhancement and geometric integrity.

9.

SEM images of the stent structure: (a) and (b) before electropolishing; (c) and (d) after electropolishing.

Figure Quantification of stent dimensional evolution before and after electropolishing via axial view analysis. Results indicate: in the as-printed L-PBF state, the stent wall thickness reached 220 μm, representing a significant 46.7% increase over the design target (150 μm); following electropolishing, the wall thickness was reduced to 145 μm, narrowing the dimensional deviation to 3.3% and validating the efficacy of this process for precision control. Notably, systematic geometric distortion was observed in the stent units irrespective of polishing: straight connecting struts exhibited degradation from the designed prismatic shape to rounded cylinders, while curved struts showed severe dimensional deviations (Figure ). Mechanistic studies revealed this distortion stems from the synergistic coupling of three factors: , Deflection-scale effect: Excessive deflection in curved struts, combined with the microscale structure (feature size 150 μm), caused laser scan path deviation. This led to asymmetric melt track overlapping and material accumulation. Process-geometry conflict: When structural dimensions (150 μm), laser spot size (50 μm), and powder particle size (15–53 μm) are of comparable magnitude, the curvature radius of the melt pool solid–liquid interface becomes comparable to the structural size. This enables surface tension-driven balling energy to dominate over gravitational energy, forcing the cross-section toward a rounded profile. Dynamic melting-solidification defects: Although the curved strut design avoided horizontal overhangs, its high deflection characteristic intensified melt material superposition effects. This increased the actual melt-solidified area by 35–40%, concurrently superimposing solidification shrinkage stress and further amplifying dimensional deviations. Such distortions are detrimental to the flexibility of vascular stents, constraining their implantation functionality. Therefore, subsequent structural designs should account for the limitations of L-PBF process parameters, particularly avoiding structures with excessive deflection in microscale components.

10.

Axial views of the vascular stent: (a) before electropolishing; (b) after electropolishing.

3.2. Tensile Properties

Figure presents the typical engineering stress–strain curve of the tensile specimen fabricated using identical processing parameters as the vascular stent. The material exhibited a distinct elastic deformation stage, with a measured Young’s modulus of 150 GPa. Upon entering the plastic deformation region, the yield strength was determined to be 232 MPa using the 0.2% offset method, while the ultimate tensile strength reached 385 MPa. The plastic deformation stage demonstrated significant work-hardening behavior, indicating the material’s continuous enhancement in resistance to further deformation during straining. The specimen ultimately underwent ductile fracture, with a measured fracture elongation of 16.44%.

11.

Representative engineering stress–strain curve of the microtensile specimen fabricated using identical L-PBF parameters to the vascular stent.

Table summarizes the mechanical properties of the specimens fabricated via L-PBF technology and compares them with those of wrought 316L alloy of identical chemical composition. The results indicate that the mechanical properties of the L-PBF-processed material are generally slightly inferior to those of the wrought alloy. Nevertheless, its yield strength reaches 74.1% of the value measured for the wrought alloy. This discrepancy can be attributed to the combined effects of the microstructure (particularly porosity) inherent in the L-PBF process and potential size effects. Such findings warrant further investigation to better understand the underlying phenomena and to explore potential applications in fields like orthopedic load-bearing implants, where a relatively lower yield strength might also be advantageous. However, experiments showed that the tensile specimens produced by L-PBF had poor elongation. Therefore, the stent structure was designed in the hope that it could expand successfully without breaking.

2. Comparison of Mechanical Properties between Tensile Specimens Produced by L-PBF and Wrought 316L Alloy.

	L-PBF 316L	Wrought 316L
E (GPa)	150	207
YS (MPa)	232	313
UTS (MPa)	385	657
δ (%)	16.44	53.5

3.3. Expansion Performance

Figure illustrates the morphological evolution of the stent under varying balloon inflation pressures. Measurement data revealed an initial stent outer diameter of 2.5 mm and a length of 19.5 mm. This initial dimension exceeded the design value of the 3D model, primarily attributable to the limitations imposed by laser spot diameter and powder particle size on printing accuracy during the L-PBF forming process.

12.

Stent expansion from 0 to 8 bar.

Nonuniform deformation was observed during expansion: At low pressures (<1.5 bar), the proximal and distal segments of the stent expanded first and at a faster rate, while the midbody outer diameter remained unchanged (Figure ). This phenomenon stemmed from the balloon length exceeding the stent length, causing initial stress concentration at the stent ends upon pressurization. As pressure increased (1.5 to 4 bar), the midbody segment began to expand, and the diameter discrepancy between the ends and midbody progressively diminished. Around 4 bar pressure, the support rings at both ends approached full expansion, exhibiting a reduced expansion rate, while the midbody expansion rate correspondingly accelerated.

Notably, at approximately 7 bar, a transient decrease in outer diameter occurred at both stent ends. Analysis suggests that the delayed expansion of the midbody segment relative to the ends resulted in slight warping of the ends at high pressure, causing them to detach from the balloon surface and consequently reducing the measured outer diameter. With continued pressure increase (>7 bar), accelerated expansion of the midbody segment restored stent conformity to the balloon, leading to a subsequent recovery of the end diameters.

At the maximum pressure of 8 bar, the midbody outer diameter reached 5.0 mm, while the end diameters measured 5.5 mm, and the stent length was reduced to 15.1 mm. As depicted in Figure , the final expansion ratios for both the ends and midbody segments exceeded 200% of the initial outer diameter (i.e., more than doubled). Microscopic images of the stent connections before and after expansion, provided in Figure , further corroborate its excellent structural integrity. No cracks were observed upon expansion to 5.0 mm, indicating superior expansion performance.

13.

Expansion rate of the stent outer diameter.

14.

Microscopic images of the stent junction: (a), (c) before expansion; (b), (d) after expansion.

In summary, a well-designed vascular stent structure enables successful expansion under single, noncyclic loading conditions. However, its expansion behavior under cyclic loading (e.g., fatigue testing) remains unclear and requires further validation.

3.4. Axial Shortening and Radial Recoil

The radial recoil rate and axial shortening rate of the vascular stent were determined through expansion experiments. Preliminary experiments revealed that employing a balloon longer than the stent length resulted in nonuniform expansion (Figure ). To optimize expansion uniformity, the balloon catheter was switched to model PB-20E-0420 (maximum expansion diameter: 4 mm, working length: 20 mm), whose working length matched the stent length, effectively reducing expansion heterogeneity. Following the expansion of the M-type stent from an initial diameter of 2.5 mm to the target diameter of 3.6 mm (Figure ), the following phenomena were observed: As pressure increased incrementally, the stent outer diameter expanded steadily without circumferential twisting, and the strut connector structures remained intact. Upon balloon deflation, the stent underwent radial contraction due to elastic recovery. To quantify the radial recoil rate and axial shortening rate after full expansion, the outer diameter and length of the stent were measured using microscopy under three distinct states, as summarized in Table .

15.

Deformation states of the M-type stent.

3. Dimensions of the Stent under Different Experimental States (Unit: mm).

M-type	D initial	L initial
As-fabricated	2.54	19.58
Fully expanded	3.63	18.41
Unloaded	3.58	18.49

By substituting the data from Table into eq and , the radial recoil rate and axial shortening rate of the M-type stent were calculated as 1.37% and 5.56%, respectively. The low radial recoil rate (1.37%) stems from the unequal-height support ring design, which promotes greater plastic deformation in smaller support units during expansion, thereby reducing the elastic deformation zone and diminishing elastic recovery capability. Clinically, this low recoil rate (<5%) ensures tight apposition between the stent and vessel wall, prevents postimplantation migration, and satisfies the radial stability requirements of ISO 25539–2. Regarding the axial shortening performance (5.56%), the theoretical design anticipated that the M-type stent’s bending struts would compensate for support ring height reduction through axial elongation. However, experimental results demonstrate that this functional mechanism failed to materialize. Analysis attributes this to two factors: (1) Geometric constraints: High-curvature connecting struts exhibit limited effective elongation capacity; (2) manufacturing defects: In the L-PBF process, the relatively large laser spot diameter (∼40 μm) and layer thickness (20 μm) cause material accumulation in curved regions (Figure d), while the absence of support structures exacerbates dimensional errors and suppresses axial extension. Axial shortening exceeding 5% may alter stent implantation positioning and compromise vascular coverage effectiveness. Current L-PBF technology remains insufficient for achieving precision in microscale forming applications; consequently, subsequent structural designs must integrate L-PBF processing characteristics. ,

3.5. Radial Support Performance

The radial compression load–displacement curve of the M-type stent (Figure ) reveals a three-stage deformation characteristic: Initial nonlinear stage: Minor geometric nonuniformity postexpansion causes incomplete contact between the upper compression plate and the stent, resulting in nonlinear load growth with displacement; elastic deformation stage: After the plate feed reaches a critical point, full stent-plate contact establishes a linear load–displacement relationship (slope representing structural stiffness); plastic strengthening stage: Continued compression induces irreversible plastic deformation, driving nonlinear load surge. Strain hardening further accelerates the increasing rate.

16.

Load–displacement curve of the M-type vascular stent under radial compression.

Per ISO 25539–2, the radial strengthdefined as the force per unit length at 50% stent diameter reduction (eq )was calculated as 840 mN/mm, meeting implant standards. This exceptional performance originates from the unique reverse-aligned unequal-height support ring design: Postexpansion, enlarged interunit angles enhance circumferential load-bearing efficiency; smaller support units undergo prioritized plastic deformation, reinforcing structural stability; increased plastic deformation zones optimize stress distribution and delay local instability.

Clinically, the 840 mN/mm radial strength significantly exceeds commercial cobalt–chromium alloy stents (≈500–700 mN/mm), demonstrating that it: effectively resists circumferential vessel wall pressure (>300 mmHg); satisfies ISO 25539–2 requirements for permanent implants; provides reliable support for lesioned vessels, reducing postinterventional recoil risk.

Although this study successfully fabricated an M-type metallic vascular stent with favorable radial support performance and low recoil rate via L-PBF technology, geometric deviations induced by melt pool behavior remain present in high-curvature connecting regions. To further improve manufacturing quality and performance consistency, subsequent work will focus on optimizing L-PBF process parameters (e.g., laser power, scan speed, layer thickness) to minimize balling and material accumulation in microscale features; investigating the influence of heat treatment processes on the microstructural evolution and mechanical properties (expandability, radial strength, and fatigue life) of the stent; and systematically evaluating the application potential and improvement strategies of additive manufacturing technology for personalized vascular stents through comparative analysis with conventionally laser-cut M-type stents in terms of microstructure and comprehensive performance.

4. Conclusions

This study reports the additive manufacturing, postprocessing, and mechanical testing of a 316L stainless steel vascular stent featuring a novel structural topology specifically designed for L-PBF. The stent was successfully fabricated, and its expansion and radial support performance were evaluated using balloon catheter expansion and flat-plate compression. The principal findings are summarized as follows:

• 1Structural Design and Manufacturing Feasibility: The synergistic design of double-period unequal-height support rings interconnected by M-shaped curved struts effectively optimized radial force distribution (reaching 840 mN/mm), satisfying ISO 25539–2 mechanical requirements for permanent implants. L-PBF technology enabled unsupported monolithic fabrication of stent arrays with regular arrangement and no collapse, validating its capability to form intricate microfeatures (150 μm wall thickness).
• 2Surface Quality Control: As-fabricated L-PBF stents exhibited adhered sintered particles and scaled melt pool morphology. Electropolishing significantly reduced surface roughness while decreasing radial dimensional deviation from 46.7% to 3.3%. Geometric distortion occurred in curved strut regions due to melt pool dynamics (balling and material accumulation), necessitating avoidance of high-curvature features in future designs to improve manufacturing precision.
• 3Micro-Tensile Properties: The L-PBF-fabricated specimen achieved a yield strength of 232 MPa (74.1% of wrought 316L alloy) and fracture elongation of 16.44%, confirming sufficient plastic deformation capacity to mitigate fracture risks during balloon expansion.
• 4
  Stent Mechanical Performance:

  • Performance: Under a single load, the stent expanded from 2.5 mm to 5.0 mm (>200% expansion ratio) without structural cracking. Its radial recoil rate (1.37%, <5%) was attributed to the unequal-height rings promoting plastic deformation and ensuring postimplantation wall apposition stability.
  • Radial Compression Performance: A radial strength of 840 mN/mm met implantation requirements. The reverse-aligned unequal-height design enhanced circumferential load-bearing efficiency by enlarging interunit angles and optimizing plastic deformation zone distribution.

Axial Shortening Limitation: The axial shortening rate (5.56%) exceeded expectations, primarily due to manufacturing-induced distortion in M-shaped struts suppressing their compensatory elongation mechanism.

The M-type stent’s high radial strength and low recoil offer a new strategy for personalized implants. L-PBF combined with electropolishing provides a viable pathway for fabricating customized devices (e.g., bifurcated stents). To address strut distortion limiting flexibility, future work should integrate topology optimization (e.g., reducing curvature) with microprocess refinement (e.g., reduced spot size/layer thickness). Subsequent studies must include fatigue life testing, hemodynamic simulations, and in vivo biological evaluation.

The authors declare no competing financial interest.