Mechanically governed fracture resistance of biodegradable Zn-Cu alloy stents: Role of ring-length optimized structural design

Abstract

Zinc alloys containing copper, featuring highly soluble biocompatible elements, are widely regarded as one of the most promising candidates for use in biodegradable stents. However, suboptimal structural designs in biodegradable zinc alloy stents often lead to immediate strut fracture upon deployment. Although ring length plays a critical role in this failure mechanism, systematic studies focusing on enhancing fracture resistance through structural design-rather than material modification-remain limited. This study engineered three distinct ring length configurations (approximating radial strengths of 89 kPa, 120 kPa, and 150 kPa) to elucidate structural optimization effects on fracture resistance during biodegradable zinc alloy stents expansion. Our results demonstrate that stents with 89 kPa and 120 kPa radial strength exhibit superior fracture resistance, whereas the 150 kPa design shows significantly elevated fracture incidence. Mechanistic analyses reveal that the capacity for geometric plasticity accommodation constitutes the dominant fracture-resistant mechanism, beyond intrinsic material properties. This capacity is achieved through stress redistribution, which mitigates localized peak stress. Optimized stents achieved uniform expansion, perfect vessel apposition, and preserved structural continuity. Histological analysis revealed a confluent endothelial layer covering the stent struts at 1 month. These findings reveal a direct relationship between structural plasticity accommodation capacity and mechanical integrity preservation, providing critical insights for developing next-generation bioresorbable stents with enhanced structural reliability.

Graphical abstract

Highlights

• •Ring length serves as a critical design parameter governing fracture resistance in biodegradable stents.
• •Shorter ring unit length (1.30 vs 1.45 mm) increases radial strength 69% but exacerbates stress concentration at crown apices.
• •Stress redistribution to low-strain zones prevents fracture despite identical Zn-1.0Cu composition.
• •Optimized stent (120 kPa) achieves complete endothelialization and perfect apposition within 1 month in vivo.

1. Introduction

Coronary artery disease (CAD) is a leading cause of global mortality and morbidity [1,2], with percutaneous coronary intervention (PCI) using stents emerging as the preferred revascularization approach due to its minimally invasive nature and reduced perioperative risks compared to coronary artery bypass grafting [3]. As the first clinically implemented devices, bare metal stents (BMS) provided adequate radial support but suffered from suboptimal deliverability and alarmingly high in-stent restenosis (ISR) rates exceeding 20% [[3], [4], [5]]. These limitations were primarily attributed to their rudimentary design architecture, particularly the excessive strut thickness (132-140 μm) that exacerbated vascular injury and flow disturbances [[6], [7], [8]]. Subsequent decade-long advancements in alloy technology, notably the introduction of cobalt-chromium and platinum-chromium alloys, enabled dramatic strut thickness reduction to 80-90 μm [3]. This engineering breakthrough significantly diminished hemodynamic perturbations at the stent-tissue interface. These improvements culminated in the development of drug-eluting stents (DES), which have largely superseded BMS in clinical practice due to enhanced deliverability, superior biomechanical performance, and improved intermediate-term outcomes [[6], [7], [8], [9]]. Nevertheless, persistent challenges plague conventional permanent metallic DES. Their persistent presence as foreign bodies induces chronic vascular inflammation, predisposes to late-stage restenosis, and creates an enduring risk of thrombogenesis. These limitations have generated significant clinical interest in next-generation bioresorbable stent (BRS) [[10], [11], [12], [13], [14], [15], [16], [17]].

BRS are broadly classified into two categories: degradable polymeric stents and absorbable metallic stents [[17], [18], [19], [20], [21], [22]]. The polymeric paradigm has been dominated by poly(l-lactic acid) (PLLA)-based systems, exemplified by the Absorb Bioresorbable Vascular Scaffold (BVS) (Abbott Vascular) which received clinical approval in 2015 [23]. However, this first-generation device exhibited inherent limitations including excessive strut thickness (157 μm) and elevated vessel wall surface coverage (27%) compared to conventional cobalt-chromium alloy stents (<15%). These design shortcomings correlated with increased 3-year stent thrombosis (ST) rates (2.5% vs. 0.6% for contemporary DES platforms), ultimately prompting Abbott Vascular to discontinue the product in 2017 [24,25].

Over the past decade, biodegradable zinc alloy stents have emerged as promising alternatives, offering an optimal combination of degradation kinetics (aligned with 3-6 month vascular remodeling timelines), biocompatibility, and tunable mechanical properties (e.g., toughness-to-stiffness balance) alongside favorable manufacturability [[26], [27], [28]]. Zinc alloys containing copper, featuring highly soluble biocompatible elements, are widely regarded as one of the most promising candidates for bioresorbable implants, with Zn-1.0Cu showing significant potential for biomedical applications [[29], [30], [31], [32], [33], [34]]. Recent advancements demonstrate that copper alloying (≤1 wt%) significantly enhances the tensile performance of zinc matrices without inducing discernible secondary phase formation [[32], [33], [34]]. Compared to conventional 316L stainless steel stents, zinc-copper alloy systems exhibit superior clinical performance metrics, including enhanced trackability, reduced radial recoil (4.3% vs. 7.6%), and attenuated vascular pulsatile stress - critical attributes for restoring physiological arterial compliance during coronary interventions [35,36]. Clinical evidence further confirms the safety profile of copper alloying, showing no elevated implantation-related risks [35,36]. However, biodegradable zinc alloy stents exhibit pronounced strength-plasticity sensitivity at ambient temperature. During implantation procedures involving rapid deformation (e.g., balloon expansion), this intrinsic property may induce strut fracture due to excessive strain accumulation [37,38]. While stent architecture critically governs these failure mechanisms, systematic investigations remain scarce.

Structural modulation emerges as a pivotal strategy, given that stent design dictates clinical performance [8,[39], [40], [41], [42]]. Computational evidence indicates that finite element analysis (FEA)-driven modeling enables novel stent geometries that minimize stress concentration while preserving deployment integrity [43,44]. Our previous studies have demonstrated that ring length plays a critical role in modulating stent plastic deformation and compression resistance, with shorter ring lengths significantly enhancing radial strength [41,42]. Engineering architectures that accommodate plastic deformation through spatial stress redistribution can substantially reduce fracture propensity. Although stent structure largely governs these failure mechanisms, systematic investigations into enhancing fracture resistance through structural design-rather than material modification-remain scarce. As a novel strategy to address the limited plasticity of Zn alloys, this study aims to systematically investigate the regulatory effect of ring length, offering important implications for clinical applications.

Using a structural framework with standardized 90 μm strut thickness, we engineered three distinct biodegradable zinc alloy stents configurations exhibiting radial strengths of 89 kPa, 120 kPa (∼120 kPa for the Xience DES, a clinical gold standard), and 150 kPa. Post-compression and expansion analyses revealed that stents with ≤120 kPa radial strength maintained structural integrity, whereas those exceeding 120 kPa exhibited significantly increased fracture incidence. Animal studies confirmed that optimized stents achieved uniform expansion, perfect vessel apposition, and preserved structural continuity. Excellent endothelialization was observed at 1-month implantation. These findings establish a direct correlation between structural plasticity accommodation capacity and mechanical integrity preservation, providing critical insights for developing next-generation BRS with enhanced structural reliability.

2. Materials and methods

2.1. Material characterization

The Zn-1.0Cu (wt.%) alloy, developed by Farshine Medical Technology Co., Ltd (China), was selected for its optimized biocompatibility and mechanical properties. The extruded alloy exhibited a Poisson's ratio of 0.29, with density and Young's modulus measured as 7.1 × 10³ kg/m³ and 105.0 GPa, respectively. Uniaxial tensile testing was performed on alloy microtubes to obtain stress-strain curves (Table 1). The full stress-strain curve of the Zn-1.0Cu alloy microtubing is presented in Fig. S1.

Table 1.

The Zn alloy's true stress-strain data.

Stress (MPa)	Plastic strain
64.15	0
201.58	0.0020
258.05	0.0040
449.67	0.0247
482.08	0.0463

2.2. Dimensional and microstructural characterization

The macroscopic dimensions of the Zn–1.0Cu alloy coronary stents were measured using deep-field optical microscopy (VHX-7000, Keyence, Japan). Microstructural characterization was performed using a scanning electron microscope (SEM, Zeiss Merlin, Germany) equipped with an energy-dispersive X-ray spectroscopy (EDS) detector. All SEM observations were conducted at an accelerating voltage of 15 kV and a working distance of 8–10 mm. EDS elemental mapping and point-scan analysis were employed to examine the distribution of Cu and to verify the alloy composition.

2.3. Stent design and finite element modeling

The stent architecture in this study adopted a structural design analogous to the Magmaris® BRS (Biotronik AG), featuring S-shaped links that demonstrate enhanced bending compliance as documented in previous studies [45,46]. Fig. 1 illustrate the geometric configuration of the investigated stent. The key geometric parameters of the stent designs was shown in Fig. S2. Rectangular struts with 90 μm thickness - conforming to current industry standards (74–90 μm range) - were modeled using 8-node linear brick elements with reduced integration (C3D8R). Given the negligible thickness-to-radius ratio of the balloon catheter, it was simulated as a membrane element with reduced integration (M3D4R).

Fig. 1.

A stent structural design and parameters used in simulations; (a) Oblique view and (b) an along axial direction planar view of stent.

The stent was crimped to an outer diameter of 1.3 mm, followed by constraint removal to allow elastic recoil of the crimped configuration. Post-dilation nominal inner diameter was set to 3.0 mm, with subsequent balloon deflation enabling unconstrained stent recoil for dimensional recovery analysis. A general contact algorithm (explicit formulation) governed interfacial interactions, employing hard contact in the normal direction and a tangential friction coefficient of 0.2. To optimize computational efficiency, FEA were conducted using the Abaqus/Standard implicit solver (version 6.14-1) for quasi-static simulations, while dynamic processes were addressed through Abaqus/Explicit.

2.4. Stent parametric configuration

Radial support characteristics were modulated by varying the circumferential unit length (L_unit) of stent rings. For comparative analysis, baseline configurations with L_unit values of 1.45, 1.34, and 1.30 mm were designated as Stent-1.45, Stent-1.34, and Stent-1.30, respectively. Strut thickness was maintained at 90 μm across all designs. To achieve comparable device lengths, yielding total lengths of 9.5, 9.2, and 9.8 mm after dimensional standardization.

Surface coverage was calculated as:

$$\text{Coverage}=\frac{S_{stent}}{S_0}\ast 100\%\text{.}$$ (1)

S_Stent and S₀ correspond to stent surface area and nominal cylindrical surface area (matching stent dimensions).

2.5. Structural integrity assessment of the Zn-Cu stent

Zn-Cu alloy microtubes (outer diameter: 2.0 ± 0.01 mm) underwent precision laser machining using a femtosecond laser cutting system. Post-processing included chemical polishing and ultrasonic cleaning to achieve a final strut thickness of 90 μm. Stents were crimped onto balloons using a CX stent crimper (Blockwise, USA) and expanded at 8 atm. Morphological evolution during expansion/recoil was monitored via deep-field optical microscopy (VHX-7000, Keyence, Japan). Post-processing integrity (including crimping, expansion, and surface topography) was evaluated using deep-field optical microscopy (VHX-7000, Keyence, Japan). Radial recoil (RR) and foreshortening (FS) were quantified per ISO 25539-2 guidelines using dimensional profiles acquired during balloon deflation.

2.6. Radial strength evaluation

Radial compressive strength of expanded stents was characterized following ASTM F3067-14 [47] using a TTR2 radial force tester (Blockwise, USA) at 0.1 mm/s compression rate. Radial strength (RS, kPa) was defined as the plateau pressure at 10% compression of the nominal stent diameter (Fig. 2).

Fig. 2.

Radial strength of the stent. RS is defined as the strength of 10% compression of the initial compression diameter (x₀).

2.7. Animal model and surgical procedures

Animal experiments were conducted using Shanghai white pigs. Stents were delivered via a balloon angioplasty catheter delivery system and implanted through a percutaneous coronary intervention approach via the femoral artery. Under digital subtraction angiography (DSA; INNOVA 2100, GE, USA) guidance, zinc alloy stents (n = 3 per group) were deployed into suitable segments of the left anterior descending artery (LAD), left circumflex artery (LCX), or right coronary artery (RCA) through the femoral artery. Target vessels with a reference lumen diameter of 2.5–3.0 mm were selected for stent implantation. Balloons were inflated according to the target vessel diameter, and stents were expanded at a stent-to-artery diameter ratio of 1.1:1. The corresponding inflation pressures were chosen based on Table S1. All implantation procedures were performed by the same experienced research team. Only one stent was placed in each artery.

The experiments were carried out in a certified commercial testing laboratory (Hunan Geruosi Medical Technology Co., Ltd., China). Beginning three days after surgery, animals received daily intramuscular injections of antibiotics to prevent infection. In addition, long-term oral administration of clopidogrel sulfate (75 mg/day) and aspirin (100 mg/day) was maintained until the endpoint of specimen collection and follow-up. All animal handling procedures complied with the protocol approved by the Institutional Animal Care and Use Committee of the company, ensuring humane treatment throughout the study.

2.8. Histological analysis

All stented vascular segments were fixed in formalin solution at the 1M follow-up and then subjected to gradient dehydration using 70%, 80%, 90%, and 100% ethanol sequentially (each step for 24 h). After air-drying, samples were embedded in T7200 resin. The embedded blocks were sectioned into approximately 20 μm thick slices using a hard tissue cutting and grinding system (EXAKT E300CP & E400CS, Germany). Sections were stained with hematoxylin and eosin (H&E), and pathological assessments—including injury, inflammation, granuloma formation, and fibrosis—were performed according to established criteria reported previously [48,49]. For each section, the lumen area and the area within the internal elastic lamina (IEL) were measured, and the neointimal area was calculated as:

Neointimal area = Area within IEL−Lumen area.

2.9. Statistical analysis

Data were processed using GraphPad Prism 9.5.1. Intergroup comparisons were performed using one-way ANOVA, and multiple-group comparisons were carried out using two-way ANOVA. Data are expressed as mean ± standard deviation (SD). Statistical significance was defined as p < 0.05; “ns” denotes no significant difference.

3. Results

3.1. Microstructural characterization of the Zn-1.0Cu alloy

The microstructure and composition of the Zn–1.0Cu alloy rod are shown in Fig. S3. EDS elemental mapping (Fig. S3b) reveals that Cu is present as finely dispersed particles (green spots) within the Zn matrix (red), with no evidence of macroscopic segregation. EDS point-scan analysis (Fig. S3a) confirms a Cu content of approximately 1.1 wt%, which aligns well with the nominal composition (Zn–1.0Cu) and further verifies compositional homogeneity. According to the Zn–Cu binary phase diagram, Cu exhibits high solubility in the Zn matrix. Within the resolution limits of EDS, no distinct micron-scale second phases were detected.

3.2. Compression behavior of stents

Following FEA, Fig. 3 depict the maximum principal strain distributions of Stent-1.45, −1.34, and −1.30 during crimping and recoil processes. Fig. 3a demonstrates that the radial recoil rates after crimping and recoil were measured as 13.9%, 13.1%, and 11.9% for Stent-1.45, −1.34, and −1.30, respectively. Notably, the high-strain regions (indicated by red arrows) exhibited significant expansion with decreasing stent loop length. The maximum principal strains were predominantly localized at the crown apex of Stent-1.30 (Fig. 3).

Fig. 3.

(a) Photographs and (b) FEA showing the distribution of principal strain on Stent-1.45, Stent-1.34, and Stent-1.30 after crimping. The red arrows indicate the sites of maximum principal strain. Scale bars: 1 mm.

3.3. Integrity assessment and radial strength

In vitro balloon expansion testing revealed distinct performance characteristics among the stent designs (Fig. 4). Stent-1.45 and Stent-1.34 exhibited the most pronounced dog-boning effect at 1.5 atm, whereas Stent-1.30 showed maximal dog-boning at 2 atm. The dog-boning effect disappeared for Stent-1.45 at 2 atm and diminished gradually for Stent-1.34 at 4 atm. Under inflation conditions, Stent-1.45 and Stent-1.34 demonstrated excellent structural integrity (Fig. 4a and b). Conversely, Stent-1.30 exhibited strut fracture during expansion (Fig. 4c). The dog-boning effect was markedly reduced in Stent-1.45 and Stent-1.34 compared to Stent-1.30 (Fig. 4d). Radial recoil measurements for Stent-1.45, −1.34, and −1.30 were 4.5%, 3.0%, and 2.9%, respectively, while axial shortening rates were 4.3%, 6.2%, and 6.4%. All values complied with the requirements of ISO 25539-2 standards (<15.0%) [47]. Radial strength measurements demonstrated progressive enhancement with decreasing loop length: Stent-1.45 (89 ± 4 kPa), Stent-1.34 (120 ± 5 kPa), and Stent-1.30 (150 ± 8 kPa) (Fig. 4e). This result demonstrates that increasing structural compactness directly enhances radial load-bearing capacity.

Fig. 4.

Balloon expansion behavior of (a) Stent-1.45, (b) Stent-1.34, and (c) Stent-1.30. (d) Comparison of dog-boning, foreshortening, and recoiling effects among Stent-1.45, Stent-1.34, and Stent-1.30. (e) Measured radial strength of Stent-1.45, Stent-1.34, and Stent-1.30, compared with corresponding simulation results (n = 3). Scale bars: 1 mm.

Fig. 5 demonstrates the impact of expansion on stent integrity and the maximum principal strain distribution. Stent-1.45 and Stent-1.34 maintained structural integrity after crimping and expansion (Fig. 5a and b), whereas Stent-1.30 exhibited fractures exclusively at high-stress concentration sites (Fig. 5c shows the morphology of a representative fractured region). Statistical analysis of three parallel specimens revealed an average of 2.7 ± 0.5 fracture points per Stent-1.30, with all fractures localized to the crown regions of the stent struts.

Fig. 5.

Post-expansion integrity assessment and maximum principal strain distribution in (a) Stent-1.45, (b) Stent-1.34, and (c) Stent-1.30. Local structural fracture is observed in Stent-1.30 within high-strain regions. The inset provides a detailed view of the fracture site and its morphology. Scale bars: 1 mm.

The microscopic analysis on the fracture surfaces of failed Stent-1.30 mm samples from in vitro expansion tests, and the results are presented in Fig. 6. During crimping and expansion, the curved segments of the wavy struts (marked in the upper-right corner of Fig. 6a) serve as the primary sites of stress concentration. Fig. 6b and c shows low-magnification SEM images of the fracture surfaces corresponding to the red boxes 1 and 2, respectively. Finite element simulation (inset in Fig. 6a) indicates that the maximum principal strain is concentrated in these regions (red arrows highlight high-strain areas). The locations marked by red boxes 1 and 2 correspond precisely to the sites of maximum stress concentration predicted by FEA. Fig. 6d displays distinct lamellar cleavage steps and river patterns at higher magnification, which are characteristic of brittle fracture, indicating rapid crack propagation along grain boundaries or phase interfaces. Fig. 6e reveals numerous microvoids and dimples, suggesting localized plastic deformation accompanying fracture, although the overall fracture mode remains quasi-brittle. This microstructural appearance results directly from the synergistic effect of localized plastic deformation and brittle cleavage under stress concentration.

Fig. 6.

Fracture behavior of Zn-1.0Cu alloy coronary stents after crimping and expansion: (a) Overall SEM morphology of the stent post-crimping and expansion, with red boxes 1 and 2 marking the high-strain fracture sites; the inset presents the FEA contour plot of maximum principal strain, where red arrows highlight stress concentration locations; (b, c) Low-magnification SEM images of the fracture surfaces corresponding to red boxes 1 and 2, respectively; (d, e) High-magnification SEM images of the fracture surfaces corresponding to red boxes 1 and 2, respectively.

3.4. Structural deformation and stress distribution

Fig. 7 illustrates the structural deformation and maximum principal stress distribution throughout the crimping-expansion cycle. A progressive increase in high-stress regions was observed with reduced loop length, particularly concentrated at the crown apex of the stent struts (Fig. 7a). The results demonstrate that the proportion of low-stress regions (≤20 MPa) progressively increased for Stent-1.45, Stent-1.34, and Stent-1.30, measuring 37.21%, 32.95%, and 20.24%, respectively (Fig. 7b). Furthermore, Stent-1.30 displaying a significantly higher maximum principal stress distribution than Stent-1.45 and Stent-1.34. Specifically, the area fractions of the relative high-stress regions (≥200 kPa) for Stent-1.45, Stent-1.34, and Stent-1.30 were 8.71%, 11.31%, and 14.43%, respectively. In contrast, Stent-1.45 showed a significantly lower proportion, representing a decrease of approximately 40% compared to Stent-1.30. Considering the minimization of stent deformation and maximum principal stress alongside overall mechanical performance, Stent-1.34 demonstrated superior suitability. Consequently, Stent-1.34 was selected for the subsequent in vivo implantation studies.

Fig. 7.

Maximum principal stress distribution in (a) Stent-1.45, Stent-1.34, and Stent-1.30 following balloon inflation and elastic recoil (red arrows indicate stress concentration zones). (b) Histograms depicting the area distribution of maximum principal stress in the stents: ≤50 MPa and ≥50 MPa. Red and black arrows denote high and low values of the maximum principal stress regions, respectively.

3.5. Quantitative coronary angiography and optical coherence tomography

The in vivo evaluation, included coronary angiography and optical coherence tomography results for the Zn-1.0Cu stent immediately post-implantation and at 1 month in porcine coronary arteries (Fig. 8). Results show all zinc alloy stents were clearly visible under fluoroscopy, demonstrating good detectability during implantation. During the follow-up period, no thrombus or in-stent restenosis was detected on angiography. The Zn-1.0Cu stent was located in the left anterior descending branch (Fig. 8a–f, 8g-l), and the lumen area and diameter distributions were analysed in the distal, middle, and proximal segments of the implantation site, as well as in the 5 cm near the implantation (Fig. 8m and n). Corresponding quantitative data are summarized in Table 2. At 1 month after implantation, OCT showed that the vessel was completely endothelialised and the stent was embedded in the vessel wall (Fig. 8g–l), the stent was biocompatible with the vessel, and the signals at the edges of the stent were clear, confirming that the stent was intact and effectively supported the arterial wall. The Zn-Cu stents used here possess good inherent radiopacity. To illustrate this point, we have now provided an additional schematic comparison in the revised manuscript (see Supplementary Fig. S4): the left panel shows the stent imaged on the balloon between two radiopaque marker bands, while the right panel demonstrates the appearance at the 1-month follow-up when the vessel was not contrast-filled. This highlights that the stent can be well visualized under conditions of low vessel opacification, but may be obscured in standard DSA views depending on imaging parameters.

Fig. 8.

Quantitative coronary angiography and OCT of the Zn-1.0Cu stent in porcine coronary arteries immediately post-implantation and at 1 month in vivo. (a, b) Coronary angiograms and stent segment locations. (c, d, e, f) OCT cross-sectional images of stented segments. (g, h) The right panels show corresponding results at 1 month(i, j, k, l). Lumen area and mean diameter of the stented segments (m) immediately post-implantation and (n) at 1 month. (D, M, and P in the Fig. denote Distal, Middle, and Proximal positions, respectively). n = 4.

Table 2.

Lumen area and mean diameter of the stented segments immediately post-implantation and at 1 month.

	Project	Implant(n = 4)	1M(n = 4)
Area(mm2)	D-5cm	6.36 ± 1.05	6.81 ± 2.93
	D	5.75 ± 0.68	3.47 ± 0.93
	M	6.44 ± 1.36	4.69 ± 1.26
	P	5.97 ± 1.35	4.42 ± 21
	P-5cm	8.59 ± 1.41	7.77 ± 1.56
Mean Diameter(mm)	D-5cm	2.83 ± 0.23	2.89 ± 0.64
	D	2.70 ± 0.16	2.09 ± 0.28
	M	2.85 ± 0.33	2.42 ± 0.33
	P	2.74 ± 0.31	2.33 ± 0.50
	P-5cm	3.30 ± 0.28	3.13 ± 0.31

3.6. Micro-CT and histological analysis

To obtain detailed morphological data on stent biodegradation, high-resolution microcomputed tomography (micro-CT) was performed on arterial sections excised immediately after stent implantation and 1 month later. Fig. 9 shows the results of micro-computed tomography and hard tissue sections of Zn-1.0Cu stents 1 month after implantation in porcine coronary arteries. Micro-CT analysis showed that the stent structure was generally intact, and no obvious stent fracture was observed. After 3D reconstruction, volume percentage of the stent at implantation and 1 month after implantation showed no significant difference between the two (Fig. 9b). Corresponding histological sections (Fig. 9c) show hematoxylin and eosin (H&E) stained cross sections of the stented vessels at the follow-up time point. The stent lumen remained patent with a smooth surface and part of the endothelium covered around the stent. The stent beams were structurally intact, with no obvious inflammatory reaction, necrosis, or fibrosis, suggesting that the zinc alloy stents were well biocompatible. The stent was well covered with neointima, and the percentage of neointimal area was 31.62% ± 2.40%.

Fig. 9.

(a) Micro-CT reconstruction.No apparent degradation happened at 1 month after implantation. (b)Volume change of the stent residuals. (c) Representative low- and high-magnification photomicrographs of hematoxylin-eosin (H&E) stained sections of abdominal aorta after 1month implantation.

At the 1-month postoperative follow-up, all stented vascular segments exhibited intact stent strut architecture and complete coverage of the stent surface by regenerated endothelium. Focal aggregates of lymphocytes were observed surrounding the zinc alloy stents, whereas no significant pathological alterations—including vascular injury, necrosis, fibrosis, or thrombus formation—were detected. Histopathological scoring was performed on a scale of 0–3, and the results are summarized in Table 3 and Fig. 8c. Quantitative analysis revealed that the mean area within the internal elastic lamina, mean lumen area, and mean neointimal area of the stented vessels were 4.12471 mm², 2.84469 mm², and 1.28002 mm², respectively (Table 4).

Table 3.

Pathological injury scores of stented vascular segments.

Lesion Number	Endothelialization	Inflammatory Reaction	Thrombosis Formation	Intimal Hyperplasia	Vascular Injury	Necrosis	Fibrosi
Stent 1	3	1	0	3	0	0	0
Stent 2	3	1	0	3	0	0	0
Stent 3	3	1	0	3	0	0	0

Table 4.

Pathological measurement results of stented vascular segments.

Lesion Number	Internal Elastic Lamina Area (mm2)	Lumen Area (mm2)	Neointimal Area (mm2)
Stent 1	4.109875	2.802175	1.30770
Stent 2	4.030375	2.853475	1.17690
Stent 3	4.233875	2.878425	1.35545

3.7. Scanning electron microscopy and energy-dispersive X-ray spectroscopy

Micro-CT provided a macroscopic assessment, while cross-sectional scanning electron microscopy (SEM) provided detailed microscopic information. To characterise the degradation products of the Zn-1.0Cu stents during vascular healing, qualitative elemental analysis of the interface shown in Fig. 10 was carried out using energy dispersive X-ray spectroscopy (EDS). After 1 month, the stent-tissue interface appeared to be intact in its entirety. The layer of degradation products with a thickness of less than 1 μm was mainly composed of oxygen. Compared with the immediate implantation, the surface elements of the implant in January showed significant enrichment of P, Cl, and Ca. P and Ca elements were regionally distributed in the stent, which was more consistent, and Cl was relatively homogeneous. The enrichment of P and Ca may originate from the biomineralisation reaction triggered by the local microenvironment during the metal degradation process. The uniform distribution of Cl element suggests that the continuous penetration of body fluid ions to the surface of the stents, which participates in the electrochemical corrosion process, is conducive to the uniform degradation of the stents.

Fig. 10.

(a) SEM elemental mapping and EDS analysis of the Zn-0.1Cu stent interface (b) immediately post-implantation and (c) at 1 month.

4. Discussion

4.1. Material-structure synergistic mechanism in biodegradable zinc alloy stents

Zinc-based alloys are considered promising BRS for cardiovascular stents due to their favorable biocompatibility and degradation profile. However, zinc exhibits an exceptionally low recrystallization temperature (approximately 10 °C), leading to significant dynamic recovery during deformation. This readily promotes strain localization, potentially causing premature fracture of the stent device [50,51]. This sensitivity may result in strut fracture during implantation due to excessive deformation or strain rate-dependent plastic instability [37,38]. Such mechanical limitations present significant challenges in biodegradable zinc alloy stents development. While structural design critically influences these behaviors, systematic investigations remain scarce.

From a plasticity perspective, stent expansion involves two interdependent components: intrinsic material plasticity and structurally enabled plasticity. While material plasticity remains an inherent property determined by composition and manufacturing processes, structural plasticity offers adjustable design parameters. Non-uniform deformation patterns with localized stress concentrations that may initiate microcracks or catastrophic fractures. Fracture resistance essentially depends on the deformation accommodation capacity provided by structural design - enhanced structural plasticity reduces fracture susceptibility. As demonstrated in Table 5, these parametric relationships underscore the importance of design optimization. Residual stress concentrations at bending regions particularly predispose to structural failure [37]. Structural optimization effectively mitigates these issues, highlighting its pivotal role in advancing BRS technologies. This section elucidates the synergistic mechanism between copper-alloyed zinc matrices and ring-length optimized architectures, which collectively govern fracture resistance through enhanced structural plasticity accommodation.

Table 5.

Optimization result of biodegradable zinc alloy stents structure compared to the Stent-1.45, Stent-1.34 and Stent-1.30.

Design	Geometry structure size				Radial strength (kPa)	Radial recoiling (%)	Fore-shorting (%)	Coverage (%)
	Thickness (mm)	Lunit (mm)	Number of stent sections	W (mm)
Stent-1.45	0.09	1.45	6	0.15	89	2.7	4.3	16
Stent-1.34	0.09	1.34	6	0.15	120	1.8	6.2	16
Stent-1.30	0.09	1.30	7	0.15	150	1.1	6.4	17

A pivotal finding of this study is the established trade-off balance between material and structure, which reveals that strategic structural design can provide adaptable plasticity through intelligent allocation of deformation space, thereby reducing sole reliance on intrinsic material properties. The results demonstrate that strategic structural design can enhance deformation accommodation while maintaining requisite mechanical strength. This paradigm shift emphasizes the untapped potential of structural design in biodegradable zinc alloy stents development.

4.2. Correlation between microstructure and stent fracture resistance

Zinc alloys incorporating copper, which features highly soluble and biocompatible elements, are widely regarded as one of the most promising candidates for bioresorbable implants, with Zn-1.0Cu showing significant potential for biomedical applications [15,37,[52], [53], [54]]. Nevertheless, tensile tests reported in the literature have shown that Zn-based alloy stents are highly sensitive to deformation. Under large strain conditions, the alloy may exhibit enhanced strain hardening, which could contribute to improved fracture resistance [37,[52], [53], [54]].

Integrating the fracture observations after crimping and expansion (Fig. 5, Fig. 6), the fracture of Zn-1.0Cu alloy coronary stents is identified as a stress-concentration-driven quasi-brittle process. The fracture sites marked by red boxes 1 and 2 in Fig. 6a coincide exactly with the regions of high stress/strain concentration predicted by FEM. This precise “geometric feature → stress concentration → fracture location” correlation provides compelling evidence for a stress-concentration-dominated fracture mechanism. The role of microstructure in fracture resistance is summarized as follows:

Dual Effects of Fine Grains: Fine grains impede dislocation motion via grain boundaries, thereby increasing the alloy's strength and resistance to plastic deformation during crimping and expansion. However, in regions of high stress concentration (e.g., the bends of wavy struts), the large grain-boundary area facilitates microcrack initiation and accelerates crack propagation under high stress.

Contribution of Compositional Homogeneity: The homogeneous distribution of Cu prevents macroscopic segregation, reducing the formation of coarse brittle phases and mitigating local stress concentrations. Nevertheless, this homogeneity is insufficient to fully suppress microcrack initiation and propagation in areas subjected to severe geometric stress concentration.

Influence of Dispersed Second Phases: Although no micron-scale second phases were detected at the EDS resolution, nanoscale ε-CuZn₄ precipitates are likely present. These nanoscale precipitates contribute primarily to alloy strengthening and have a limited effect on hindering crack propagation. Instead, cracks tend to propagate rapidly along the boundaries of fine grains, constituting the dominant fracture path.

In summary, the synergistic effect of the fine-grained structure and dispersed ε-CuZn₄ precipitates in the Zn–1.0Cu alloy, combined with geometric stress concentration at the bends of wavy struts during crimping and expansion, triggers the failure sequence of “microcrack initiation → cleavage propagation → brittle fracture. The event chain during stent deformation is as follows: geometric features of wavy struts → local stress concentration → microcrack initiation at grain/phase boundaries → crack propagation along the stress concentration direction (accompanied by localized plasticity) → final formation of a fracture surface exhibiting “lamellar tearing + microvoid coalescence.” This logical pathway directly confirms the core mechanism of “stress concentration → fracture.”

4.3. Influence of structural design on stress-corrosion of biodegradable zinc-alloy stents

The standard corrosion potential of zinc (−0.76 V) lies between that of iron (−0.44 V) and magnesium (−2.37 V), contributing to a favorable corrosion rate for stent materials [37]. Nevertheless, despite the attractive degradation characteristics and biocompatibility of biodegradable zinc alloy stents, material composition (e.g., Zn-Cu systems) determines baseline properties, structural design significantly influences fracture resistance related to radial strength - an area requiring systematic investigation. From a structural perspective, achieving uniform deformation in zinc alloy stents is crucial, as it minimizes stress concentrations, thereby improving mechanical integrity and biocompatibility during service.

Guided by the post-expansion stress distribution obtained from FEA, preliminary in-vitro immersion tests were conducted to investigate stress-corrosion interactions. Stents were statically immersed in phosphate-buffered saline (PBS) at 37 °C. The results revealed distinct differences in corrosion initiation sites (Supplementary Fig. S5). In non-expanded (stress-free) stents, initial pitting occurred randomly (Supplementary Fig. S5a and b).

In contrast, stents subjected to crimping and expansion consistently exhibited corrosion initiation at specific, predetermined locations (Supplementary Fig. S5c and d). Notably, these corrosion-initiation sites colocalized precisely with the regions of maximum residual tensile stress predicted by FEA (yellow arrows indicate corrosion sites, and red arrows denote maximum-stress regions in Fig. S5d). Corrosion subsequently propagated from these high-stress initiation points to adjacent areas (Supplementary Fig. S5e and f).

These preliminary findings provide direct experimental evidence that localized stress concentration governs the onset of corrosion in Zn–Cu stents, indicating susceptibility to stress corrosion. This strongly suggests that the stent's structural design, by dictating the residual stress-field distribution, profoundly influences its corrosion behavior and likely its overall degradation mode.

It is worth noting that residual stress has been shown to accelerate corrosion-assisted fracture in biodegradable magnesium alloys [44]. Therefore, investigating the long-term impact of stress-corrosion coupling on the integrity of biodegradable zinc-alloy stents will be a key focus of future research.

4.4. Perspectives on structural design development for biodegradable zinc alloy stents

A comparison of strut thickness between the optimized stent, commercially available DES, and representative BRS made from different materials is presented in Fig. 11 [12,27,30,54]. It should be noted that magnesium-based stents, despite their greater strut thickness, exhibit rapid degradation rates and a relatively high incidence of restenosis. In contrast, iron-based stents show very slow corrosion rates and considerable accumulation of corrosion products. Although the first-in-human implantation of an iron-based drug-eluting coronary stent system was reported in 2018, further clinical evidence is required to confirm long-term safety and efficacy [37,53,54].

Fig. 11.

The strut thickness of the optimized Zn-Cu stent with commercially available DES and representative BRS made from different materials [12,27,30,52]. (1G DES:The first generation drug-eluting stents; BRS: bioresorbable stent; CoCr: cobalt-chromium; PLA: polylactic acid).

Our study addresses this knowledge gap by evaluating three structural prototypes with distinct radial strengths (89 kPa, 120 kPa, 150 kPa). The Stent-1.34 demonstrates performance metrics approaching those of the Xience DES -considered the clinical gold standard - particularly in radial strength (120 kPa) and surface coverage(16% vs. 15%) [12]. The strut thicknesses of biodegradable polymer-based and Magmaris™ stents are significantly greater than those of commonly used second-generation DES.

Although long-term data (3-6 months) are essential for comprehensive evaluation of biodegradable stents, we contend that 1-month findings remain highly valuable. This early phase captures the critical period of initial mechanical stability, acute inflammatory response, and early endothelialization—all prerequisites for long-term success. In our planned long-term studies (3–6 months and beyond), we aim to collect sufficient degraded material for analysis using XRD alongside complementary spectroscopic techniques. This will allow direct confirmation of the crystallographic phases of degradation products, enabling us to elucidate the detailed biodegradation mechanism and correlate it with vascular remodeling outcomes.

Table 6 compares key performance metrics of the Zn-1.0Cu stent with commercially available benchmarks (Absorb GT1 and Xience). As summarized in Table 6, the Zn-1.0Cu stent exhibits properties comparable to those of Xience V DES-considered the clinical gold standard-particularly in terms of radial strength (120 kPa), acute recoil (1.8% vs. 3.6%), and surface coverage (16% vs. 15%). The combination of thin struts, low acute recoil, and a tailored degradation rate that aligns with the 3-6 month vascular remodeling window [28], positions zinc-based alloys as promising candidates for next-generation BRS. These collective advantages in mechanical and degradation performance provide a theoretical foundation for the development of future BRS.

Table 6.

Comprehensive comparison of biodegradable zinc alloy stents, Absorb GT1and Xience stents [12,27,30,45].

Φ3.0 mm	Strut thickness [μm]	Crossing profile [mm]	Recoil [%]	Radial strength [kPa]	Coverage [%]
Pure Zn stent [30]	165	-	-	41.06	-
Zn–0.1Li stent [30]	65	1.09	0.47	82.42	-
Zn–0.8Cu stent [27]	100	0.8-1	4.0	114	-
Zn-1.0Cu	90	1.20	1.8	120	16
Absorb GT1 [45]	150	1.38	5.22	120	27
Xience Prime [12]	81	1.11	3.6	120	15

Our findings on the relationship between radial strength and structural design establish a framework for optimizing biodegradable zinc alloy stents development. Notwithstanding the limitations of this study, such as sample size constraints and the absence of long-term implantation data, the results offer valuable insights for optimizing stent design to enhance fracture resistance in biodegradable zinc-copper alloy stents. The results demonstrate that structural design represents a viable approach not only to improve mechanical performance but also to mitigate potential failure modes by reinforcing the overall mechanical integrity of the alloy. The established optimization framework and fundamental structure-property relationships lay the groundwork for future advancements in BRS technology.

Our systematic investigation of structural designs provides valuable insights for biodegradable zinc alloy stents development. Future research should focus on three key directions: (1) Computational optimization of geometric parameters using machine learning algorithms, (2) Multiscale investigation of degradation-structure-mechanical property relationships, and (3) Development of hybrid material-structure design frameworks. This study establishes a foundational framework for balancing biomechanical performance and fracture resistance in biodegradable zinc alloy stents, accelerating their clinical translation.

5. Conclusions

Based on the findings of this study, the following conclusions are drawn:

• (1)Radial Strength-Structure Trade-off: Ring length serves as a critical design parameter governing fracture resistance in biodegradable stents. Reducing ring unit length from 1.45 mm to 1.30 mm enhanced radial strength by 69% (89 to 150 kPa), but exacerbated stress concentration at crown apices, increasing high-stress regions by 66% (8.71% to 14.43% area fraction ≥200 MPa).
• (2)Fracture Resistance Mechanism: Stents with ≤120 kPa radial strength (Stent-1.45/1.34) preserved structural integrity by redistributing stress to low-strain regions (≤20 MPa area: 37.21% vs. 32.95% vs. 20.24% in Stent-1.30). This geometric plasticity accommodation suppressed fracture initiation despite identical material properties.
• (3)Clinical Translation Potential: The optimized Zn-Cu stent achieved homogeneous expansion, perfect vessel apposition, and complete endothelialization within 1 month, confirming biocompatibility and structural continuity in vivo.

These insights establish that fracture resistance is governed by structure, providing a framework for balancing biomechanical performance and biological response in next-generation BRS, and accelerating the clinical translation of Zn-based implants.

CRediT authorship contribution statement

Yafei Li: Writing – original draft. Lei Wang: Data curation. Jie Li: Data curation. Penghui Zeng: Data curation. Shaokang Guan: Writing – review & editing. Hongtao Yang: Writing – review & editing. Jiang Liu: Project administration, Funding acquisition. Yufeng Zheng: Resources, Project administration.

Ethics approval and consent to participate

These experiments were conducted at a commercial testing laboratory (Shenzhen Advanced Medical Services Co., Ltd). Animals were humanely treated in accordance with protocols approved by the institutional animal care and use committee for the company (Approval number: IACUC20250801).

Declaration of competing interest

The authors declare the following financial interests/personal relationships which may be considered as potential competing interests: Yufeng Zheng is an editor-in-chief and Shaokang Guan is an editorial board members for Bioactive Materials and was not involved in the editorial review or the decision to publish this article. Yafei Li, Lei Wang, Jie Li, Penghui Zeng and Jiang Liu are currently employed by Farshine Medical Technology Co. Limited.

Other authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

Appendix A. Supplementary data

The following is the Supplementary data to this article: