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<figure><img src="image_86.png" /></figure>

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OPEN ACCESS

EDITED BY
Yawei Du,
Shanghai Jiao Tong University, China

REVIEWED BY
Junjie Li,
Soochow University, China
Weifeng Zeng,
University of Wisconsin-Madison,
United States

*CORRESPONDENCE
Yuan-Yu Hsueh,
yyhsueh@mail.ncku.edu.tw

RECEIVED 04 May 2023
ACCEPTED 23 May 2023
PUBLISHED 30 May 2023

CITATION
Hsu M-W, Chen S-H, Tseng W-L, Hung K-S, Chung T-C, Lin S-C, Koo J and Hsueh Y-Y (2023), Physical processing for decellularized nerve xenograft in peripheral nerve regeneration.
*Front. Bioeng. Biotechnol.* 11:1217067.
doi: 10.3389/fbioe.2023.1217067

COPYRIGHT
© 2023 Hsu, Chen, Tseng, Hung, Chung, Lin, Koo and Hsueh. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.

# Physical processing for decellularized nerve xenograft in peripheral nerve regeneration

Ming-Wei Hsu¹, Szu-Han Chen¹,²,³, Wan-Ling Tseng⁴, Kuo-Shu Hung¹, Tzu-Chun Chung⁵, Sheng-Che Lin⁶, Jahyun Koo⁷ and Yuan-Yu Hsueh¹,²,³*

¹Division of Plastic and Reconstructive Surgery, Department of Surgery, National Cheng Kung University Hospital, College of Medicine, National Cheng Kung University, Tainan, Taiwan, ²Center of Cell Therapy, National Cheng Kung University Hospital, College of Medicine, National Cheng Kung University, Tainan, Taiwan, ³International Research Center for Wound Repair and Regeneration, National Cheng Kung University, Tainan, Taiwan, ⁴Division of Plastic and Reconstructive Surgery, Department of Surgery, Tainan Hospital, Ministry of Health and Welfare, Tainan, Taiwan, ⁵Department of Orthopedic Surgery, E-Da Hospital, Kaohsiung, Taiwan, ⁶Division of Plastic Surgery, Department of Surgery, An-Nan Hospital, China Medical University, Tainan, Taiwan, ⁷School of Biomedical Engineering, Korea University, Seoul, Republic of Korea

In severe or complex cases of peripheral nerve injuries, autologous nerve grafts are the gold standard yielding promising results, but limited availability and donor site morbidity are some of its disadvantages. Although biological or synthetic substitutes are commonly used, clinical outcomes are inconsistent. Biomimetic alternatives derived from allogenic or xenogenic sources offer an attractive off-the-shelf supply, and the key to successful peripheral nerve regeneration focuses on an effective decellularization process. In addition to chemical and enzymatic decellularization protocols, physical processes might offer identical efficiency. In this comprehensive minireview, we summarize recent advances in the physical methods for decellularized nerve xenograft, focusing on the effects of cellular debris clearance and stability of the native architecture of a xenograft. Furthermore, we compare and summarize the advantages and disadvantages, indicating the future challenges and opportunities in developing multidisciplinary processes for decellularized nerve xenograft.

KEYWORDS

decellularized nerve xenograft, peripheral nerve regeneration, physical processing, freeze-thaw, perfusion, immersion and agitation, sonication, supercritical fluids

## 1 Introduction

The current gold standard for peripheral nerve repair in segmental defects is autologous nerve grafting (Kalomiri et al., 1994). However, several limitations hamper the clinical practice, such as limited supply, donor site morbidity, and size discrepancy (Hu et al., 2009). Alternative autologous nerve substitutes, such as autologous veins, muscles, and tendons have been utilized with variable outcomes (Ray et al., 2011). Because of the abovementioned drawbacks, tissue engineered nerve grafts (TENG) were developed, aiming to provide a biomimetic scaffold for peripheral nerve regeneration (Houschyar et al., 2016). Although the Food and Drug Administration had approved several off-the-shelf synthetic nerve conduits on the market, but most of them are limited to nerve gaps <3 cm or <0.5 cm in small or large diameter nerves, respectively (de Ruiter et al., 2009). In addition, a recent review does not support the use of currently available TENG over standard nerve repair (Thomson et al.,
<figure><img src="image_87.png" /><figcaption>FIGURE 1</figcaption></figure>

The physical processing methods for tissue decellularization of nerve xenograft, including freezing and thawing cycle, sonication, immersion and agitation, perfusion and supercritical fluids.

2022). To achieve a better outcome in longer nerve gaps, nerve allografts are harvested and processed under a commercial decellularization process. Avance®, a mature commercial product which is produced by processing human nerve tissue with a combination of detergent decellularization, chondroitinase CSPG degradation, and gamma-irradiation sterilization, has clinical evidence to overcome up to 70 mm nerve gap in sensory, mixed, and motor nerve repair (Gunn et al., 2010; Safa et al., 2020). However, limited donor sources, high costs, low temperature preservation, and potential immune rejection still remain as clinical concerns for its wide usage.

Xenotransplantation was developed due to the unlimited availability of sources. In general, fresh xenografts elicit an immune response that causes graft rejection. The presence of non-self-antigenic epitopes triggers the activation of T- and B-lymphocytes. This, in turn, leads to the activation of an immune response mediated by antibodies, ultimately resulting in

the rejection of these cells. (Fox et al., 2001; Vadori and Cozzi, 2014; Lopresti et al., 2015). Despite using immunosuppressive drugs, decellularization is an effective approach to utilize the xenogeneic and allogeneic tissues. The cellular components of tissues are the main cause of an adverse host response. To mitigate immune rejection, several decellularization techniques were developed to remove cellular components to reduce immunogenic reactions while preserving native scaffold or extracellular matrix (ECM) microstructure (Rana et al., 2017). Several recent studies reported that acellular nerve xenografts have similar effects on regeneration and immunocompatibility compared with acellular nerve allografts (Zhang et al., 2010; Huang et al., 2015). The concept of Decellularized extracellular matrix (dECM) is established as using various methods (physical, enzymatic, or chemical) to lyse cells and remove the intracellular components from a tissue while preserving the native extracellular components and the cues for cell proliferation and differentiation (Dahl et al., 2003; Lopresti et al.,
<table><caption>TABLE 1 Summary of current physical processing methods for decellularized nerve xenograft.</caption><thead><tr><td>Method</td><td>Mechanism</td><td>Xenograft source</td><td>Advantages</td><td>Disadvantages</td><td>References</td></tr></thead><tbody><tr><td>Freeze-thaw cycles</td><td>Intracellular crystallization causes cell death</td><td>porcine, rabbit, dog</td><td>easy, low-cost</td><td>limited ability to eliminate cells</td><td>Ide (1983), Ide, Tohyama, Yokota, Nitori, & Onodera (1983), Osawa et al. (1990), Evans et al. (1998), Hess et al. (2007), Jesuraj et al. (2014), Kaizawa et al. (2017), Philips et al. (2018)</td></tr><tr><td>Sonication</td><td>Rupture of the cell membrane via generating acoustic cavitation</td><td>rabbit</td><td>low-cost; can shorten processing time</td><td>can't remove DNA content</td><td>Hudson et al. (2004), Boriani et al. (2017), Bolognesi et al. (2022), Suss et al. (2022)</td></tr><tr><td>Perfusion</td><td>Pressure induced by perfusion via natural vasculature can remove cellular components</td><td>porcine</td><td>might overcome long nerve gap<br>avoid ischemic damage and central necrosis</td><td>surgical complexity, donor site morbidity and limited nerve availability</td><td>Wüthrich et al. (2020)</td></tr><tr><td>Immersion and Agitation</td><td>Isolate and remove cells from the tissue</td><td>rat</td><td>Easy, might overcome long nerve gap</td><td>Time consumable</td><td>Vasudevan et al. (2014)</td></tr><tr><td>Supercritical fluids</td><td>Uncertain</td><td>porcine</td><td>environmentally friendly, nontoxicity, low cost, disinfection</td><td>low ability of defatting</td><td>Isenschmid et al. (1995), White et al. (2006), Casali et al. (2018), Topuz et al. (2020), Wei et al. (2022)</td></tr></tbody></table>

2015; Kim et al., 2019). Although regeneration can occur in short nerve gaps regardless of an immune response, nerve growth was suppressed in long nerve gaps (Choi and Raisman 2003).

Zhang and Chen summarized several decellularization protocols along with their mechanisms and disadvantages, including physical, chemical, and enzymatic treatments in organ or tissue (Zhang et al., 2022). For nerve growth, a biomimetic neural scaffold is crucial, and thus, limited decellularization methods are applicable. Chemical-based decellularization was widely used in nerve tissues, but the major concern was the possible interruption of nerve growth by the residual chemical agents (Han et al., 2019). Enzymatic treatments provide high specificity for the removal of cellular components; however, they cannot be removed completely that might induce severe distortion of the ECM structure.

This minireview aims to focus on the latest five physical-based decellularization methods for peripheral nerve xenotransplantation (Figure 1). Table 1 summarizes the mechanisms, advantages, and disadvantages of the current physical processing methods.

## 2 Different approaches of physical processing

This minireview aims to clarify the latest four physical-based decellularization methods for peripheral nerve xenotransplantation (Figure 1). Table 1 summarizes the mechanisms, advantages, and disadvantages of the current physical processing methods.

### 2.1 Freeze-thaw cycles

Freeze-thaw cycles refer to a repetitive freeze-drying process (Zhang et al., 2022). In nerve xenografts, the aim of cold preservation and freeze-thaw cycles is to destroy the nerve cell membranes by inducing the formation of intracellular ice crystals, thereby reducing

the immunogenicity of nerve xenografts (Hare et al., 1993; Fox et al., 2005; Philips et al., 2018). Cold preservation and freeze-thaw cycles are easy to manipulate and an initial step in many decellularization protocols, as the protocols can be adjusted according to nerve length and diameter, depending on the laboratory preferences (Lasso and Deleyto, 2017).

However, studies have shown that the ultrastructure of nerves may be damaged by the freeze-thaw cycles, although the mechanical properties of the nerves are preserved (Osawa et al., 1990; Evans et al., 1998). A slower recovery in rats that received frozen grafts compared with those that received fresh autografts in a 2 cm median-nerve-gap rat model using Beagle dog acellular frozen xenografts indicate that freezing leads to a barren microenvironment for nerve regeneration (Accioli De Vaconcellos et al., 1999). Furthermore, the freeze-thaw process led to nerve xenograft rejection caused by the residual cells and debris (Lu et al., 2009). Despite the easy manipulation of freeze-thaw cycles, they might be responsible for damage to the nerve microstructure and a depleted environment for nerve regeneration.

### 2.2 Perfusion

Perfusion refers to the process of introducing circulating agents through the intrinsic vascular system of organs or tissues. This technique is typically used in larger, thicker tissues or whole organs (Goh et al., 2018). Only one study has reported the use of perfusion decellularization in peripheral nerve repair. Wüthrich and Lese applied perfusion decellularization to surgically procured vascularized porcine sciatic nerves (Wüthrich et al., 2020). A 3D microcomputed tomography imaging showed preserved vasculature and the ECM component. The dissected graft contained more external connective tissues, and the measurable growth factors were detectable at low levels. These results suggest that the biological activity of the graft may be retained and could
promote nerve regeneration. An *in vitro* study revealed the potential for re-endothelialization, but no *in vivo* study has been conducted yet. These nerve scaffolds can be created for specific lesions and are becoming increasingly available, with the potential to overcome large nerve gaps.

## 2.3 Immersion and agitation

Immersion and agitation refers to the process of submerging tissues into decellularization solutions with constant mechanical agitation (Crapo et al., 2011; Zhang et al., 2022). Compared with perfusion treatment, immersion and agitation is used in processing small, fragile and thin sections of tissues without innate vascular structures (Alshaikh et al., 2019). The efficiency of this method depends on different parameters including agitation intensity, decellularization agent and tissue dimension (Duisit et al., 2018). Immersion and agitation methods of tissue decellularization have been described for a wide variety of tissues, including peripheral nerves (Hudson et al., 2004; Karabekmez et al., 2009; Vasudevan et al., 2014). However, most of the decellularization solutions were chemical, detergent, or enzymatic solutions (Keane et al., 2015). To the best of our knowledge, only Vasudevan et al. reported an immersion and agitation method with detergent-free solution in peripheral nerve field (Vasudevan et al., 2014). In their design, the nerve grafts were immersed with detergent-free solution and were cultured at 37°C with 5% CO2 for 2 weeks under constant agitation, which was performed to initiate Wallerian degeneration *in vitro* to clear axonal and myelin debris inside the nerves. In the 3.5-cm sciatic nerve transection rat model, nerve regeneration was identical to that of detergent-processed grafts, while the functional nerve regeneration was only observed in detergent-free decellularized grafts at 12 weeks. This method did not significantly affect the ECM surface structure, collagen structure and integrity, mechanical strength, and GAG content, but may cause more damages to tissues due to the limited diffusion of chemical, detergent, or enzymatic decellularization solutions by agitation (Wilson et al., 2016; Simsa et al., 2019; Zhang et al., 2022).

## 2.4 Sonication

Sonication is a method of rupturing the cell membrane by generating acoustic cavitation bubbles and inducing shear stress effect. It can assist in the penetration of agents by vibration as well as remove cellular debris. It was demonstrated that coupling freeze-thawing with sonication contributes to a cell-free and aseptic xenograft in a shorter time than applying freeze-thawing alone in a rabbit peripheral nerve model (Boriani et al., 2017). Based on these results, they further developed a new method of soaking the nerve tissues in decellularizing solutions, with combination of sonication and freeze-thaw cycles (Bolognesi et al., 2022). This new method was validated through histology and immunohistochemistry, showing its application to decellularized xenografts with similar or better results compared with the Hudson technique (Hudson et al., 2004). However, it was observed that

sonication during chemical decellularization did not remove deoxyribonucleic acid (DNA) content but only cellular debris and myelin sheaths (Suss et al., 2022). So far, no study has utilized sonication as a single process for xenograft decellularization.

## 2.5 Supercritical fluids

Supercritical fluids are fluids above their critical pressure and temperature that possess characteristics such as low viscosity and high diffusivity. Supercritical carbon dioxide (ScCO₂) is an environmentally friendly solvent that is widely used in the field of biomedicine and biomaterials due to its nontoxicity, low cost, and superior disinfection and sterilization abilities (Subramaniam et al., 1997; Casali et al., 2018). It has a critical pressure 7.38 MPa and a critical temperature 31°C. ScCO₂ has been shown to remove cells from the tissues while maintaining the ECM structure. However, the exact mechanism by which this occurs is unclear. The hypothesis that high pressure induces cell bursting, as claimed by Topuz et al., has been refuted (Isenschmid et al., 1995; White et al., 2006). It is hypothesized that ScCO₂ might induce hypoxia, which has been validated by histological and morphological analyses after successful decellurization of bovine optic nerves using ScCO₂ (Topuz et al., 2020). Wei et al. developed a porcine acellular nerve xenograft based on supercritical extraction technology and validated it in a 15-mm rat sciatic nerve model (Wei et al., 2022). A porous nerve basement membrane with a well-preserved 3D structure was observed. Low cytotoxicity was noted *in vitro*, leading to decreased immune response *in vivo*. The ScCO₂ treatment group was found to be similar to the autologous nerves in terms of regenerated nerve quality, target muscle wet weight regain, and motor function recovery. Moreover, the hybrid detergent plus ScCO₂ treatment demonstrated better outcomes in terms of decellularization and defatting compared to ScCO₂ alone (Casali et al., 2018).

## 3 Challenges and opportunities

In recent decades, tissue engineering has been applied in the field of regenerative medicine to peripheral nerve repair, relying on the three main pillars: scaffolds, cells, and growth factors (Carvalho et al., 2019). In scaffolds, preserving the biomimetic microenvironment is essential, whereas in xenografts, the removal of cells is critical to prevent subsequent immune rejection. Gilpin and Yang et al. measured four aspects of the decellularized ECM to assess the quality of decellularization: removal of cells, elimination of genetic material, preservation of the protein content, and retention of the mechanical properties (Gilpin and Yang, 2017). Carpo et al. proposed specific criteria for assessing the efficacy of cell removal: the decellularized ECM must have the following: 1) less than 50 ng double-stranded DNA per mg ECM dry weight, 2) less than 200 bp DNA fragment length, and 3) no visible nuclear material by 4',6-diamidino-2-phenylindole staining (Crapo et al., 2011).

Chemical and enzymatic approaches are the most widely applied methods for peripheral nerve decellularization, but the possible toxicity of the chemicals and destruction of the ECM proteins
are major drawbacks of these methods. Physical treatments like
modulating temperature or pressure have been effective, but it
comes with limitations because of the natural architecture of
nerves. Freeze-thawing is widely used and an important step in
all decellularization protocols. However, debris retention and
potential microstructure damage might restrict the in vivo
therapeutic effects. Perfusion decellularization was developed to
overcome long nerve gaps using vascularized porcine sciatic
nerves, but the evidence of an in vivo study is lacking.
Immersion and agitation decellularization is commonly used but
generally with chemical, detergent, and/or enzymatic solutions.
Sonication is integrated with other methods to assist in removing
cellular debris. However, the clearance of DNA content is
questionable. Supercritical fluids show promise with advantages
of nontoxicity, low cost, superior disinfection and sterilization
abilities, and can solely and effectively remove cells. Pure physical
treatment has a limited but acceptable effect in decellularization
compared with chemical and enzymatic approaches. Both
immersion and agitation and supercritical fluids had shown that
identical decellularization efficacy can be achieved as compared to
chemical or enzymatic approaches. Further investigations are
required to validate the in vivo therapeutic outcomes. With
comprehensive understanding of the physical processing
methods, multidisciplinary integration of different approaches are
expected to elicit accumulative benefits, in terms of reducing
immunogenicity and preserving the mechanical properties and
microenvironment of native nerve tissue.

Author contributions

M-WH, S-HC, W-LT, K-SH, and T-CC wrote the manuscript, S-CL, JK, and Y-YH conceived the idea of the study, supervised, and revised the manuscript. All authors contributed to the article and approved the submitted version.

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Dahl, S. L. M., Koh, J., Prabhakar, V., and Niklason, L. E. (2003). Decellularized native and engineered arterial scaffolds for transplantation. *Cell Transpl.* 12 (6), 659–666.

Funding

This study was supported in part by grants from the National Science and Technology Council (111-2311-B-006-001 and 111-2923-B-006-001-MY2 to Y-YH; 111-2314-B-006-095 to S-HC) and National Cheng Kung University Hospital (NCKUH-11109014 to Y-YH) in Taiwan.

Acknowledgments

The authors would like to express their sincere gratitude to Ko-
Chung Yen and Dar-Jen Hsieh from ACRO biomedical company in
Taiwan, for the explanation of terminology and concept to the
decellularization process. The research was supported in part by
Higher Education Sprout Project, Ministry of Education to the
Headquarters of University Advancement at National Cheng
Kung University (NCKU).

Conflict of interest

The authors declare that the research was conducted in the
absence of any commercial or financial relationships that could be
construed as a potential conflict of interest.

Publisher's note

All claims expressed in this article are solely those of the authors
and do not necessarily represent those of their affiliated
organizations, or those of the publisher, the editors and the
reviewers. Any product that may be evaluated in this article, or
claim that may be made by its manufacturer, is not guaranteed or
endorsed by the publisher.

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OPEN ACCESS

EDITED BY
Yawei Du,
Shanghai Jiao Tong University, China

REVIEWED BY
Ludovica Cacopardo,
University of Pisa, Italy
Junjie Li,
Soochow University, China

*CORRESPONDENCE
Cian O'Leary,
✉ cianoleary@rcsi.ie

RECEIVED 16 March 2023
ACCEPTED 04 May 2023
PUBLISHED 06 June 2023

CITATION
Khalid T, Soriano L, Lemoine M, Cryan S-A, O'Brien FJ and O'Leary C (2023), Development of tissue-engineered tracheal scaffold with refined mechanical properties and vascularisation for tracheal regeneration. *Front. Bioeng. Biotechnol.* 11:1187500. doi: 10.3389/fbioe.2023.1187500

COPYRIGHT
© 2023 Khalid, Soriano, Lemoine, Cryan, O'Brien and O'Leary. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.

# Development of tissue-engineered tracheal scaffold with refined mechanical properties and vascularisation for tracheal regeneration

Tehreem Khalid<sup>1,2,3</sup>, Luis Soriano<sup>1,2,4</sup>, Mark Lemoine<sup>2,3,5</sup>, Sally-Ann Cryan<sup>1,2,3,4,5</sup>, Fergal J. O'Brien<sup>2,3,4,5</sup> and Cian O'Leary<sup>1,2,3*</sup>

<sup>1</sup>School of Pharmacy and Biomolecular Sciences, RCSI University of Medicine and Health Sciences, Dublin, Ireland, <sup>2</sup>Tissue Engineering Research Group, RCSI University of Medicine and Health Sciences, Dublin, Ireland, <sup>3</sup>Advanced Materials and Bioengineering Research (AMBER) Centre, RCSI and Trinity College Dublin, Dublin, Ireland, <sup>4</sup>Centre for Research in Biomedical Devices (CúRAM), NUI Galway, Galway, Ireland, <sup>5</sup>Trinity Centre for Biomedical Engineering, Trinity College Dublin, Dublin, Ireland

**Introduction:** Attempted tracheal replacement efforts thus far have had very little success. Major limiting factors have been the inability to efficiently re-vascularise and mimic the mechanical properties of native tissue. The major objective of this study was to optimise a previously developed collagen-hyaluronic acid scaffold (CHyA-B), which has shown to facilitate the growth of respiratory cells in distinct regions, as a potential tracheal replacement device.

**Methods:** A biodegradable thermoplastic polymer was 3D-printed into different designs and underwent multi-modal mechanical assessment. The 3D-printed constructs were incorporated into the CHyA-B scaffolds and subjected to *in vitro* and *ex vivo* vascularisation.

**Results:** The polymeric backbone provided sufficient strength to the CHyA-B scaffold, with yield loads of 1.31–5.17 N/mm and flexural moduli of 0.13–0.26 MPa. Angiogenic growth factor release (VEGF and bFGF) and angiogenic gene upregulation (KDR, TEK-2 and ANG-1) was detected in composite scaffolds and remained sustainable up to 14 days. Confocal microscopy and histological sectioning confirmed the presence of infiltrating blood vessel throughout composite scaffolds both *in vitro* and *ex vivo*.

**Discussion:** By addressing both the mechanical and physiological requirements of tracheal scaffolds, this work has begun to pave the way for a new therapeutic option for large tracheal defects.

KEYWORDS

trachea, tissue engineering, 3D printing, respiratory, vascularisation

## 1 Introduction

Replacement of large tracheal defects is a complex issue, with little clinical success to date in spite of other advancements in modern medicine. Two key issues facing tracheal replacement are 1) the mechanical differences between the implant and surrounding native tissue and 2) of host acceptance of implants and integration into the host
vascular network. Many promising studies have investigated novel approaches to replacing tracheal defects using additive manufacturing techniques (Les et al., 2019; Kang et al., 2019; Ricardo et al., 2019; Xia et al., 2019; Lee et al., 2020). The most notable were the Marlex™ mesh porous prosthesis, these showed promise in animal studies with 16 months of graft survival and patency (Beall et al., 1962). However, they did not perform well in humans due to stenosis and erosion of surrounding blood vessels (Maziak et al., 1996). Tracheal allografts have also been investigated since 1979 (Rose et al., 1979). Recent developments by Delaere et al. (2012), Delaere et al. (2014), and Delaere et al. (2019) saw the development of a two-step technique in which the donor trachea is revascularised in the patient's forearm before implantation. A cadaveric trachea was decellularised, implanted in the forearm of the recipient and seeded with buccal mucosa to enhance vascularisation prior to implantation (Delaere et al., 2010). This approach was further explored with four additional patients, but tracheal necrosis and poor vascularisation hampered the outcome of this clinical trial (Omori et al., 2008). 3D-printing has also been utilised to treat tracheal defects in paediatric patients as 3D printed airway splints have been granted FDA approval (Pacetti, 2020). First reported in 2013 to treat a severe case of tracheobronchomalacia (TBM) (Zopf et al., 2013), the splints have been successfully implanted in a wider set of patients with low mortalities (Les et al., 2019; Morrison1 et al., 2017; Huang et al., 2016; Morrison et al., 2017). However, their use has been restricted to paediatric patients. Only one attempt has been reported in an adult, with maintained airway patency 3-month post implantation. Overall, there is currently no established gold standard for tracheal regeneration that fully addresses all major concerns.

The mechanical discrepancy between native tracheal tissue and that of the replacement graft is often one of the major causes for implant failure in vivo and within clinical efforts (Grimmer et al., 2004; Azorin et al., 2006; Luo et al., 2013; Wood et al., 2014; Elliott et al., 2017). The trachea is composed of distinct structural regions that each fulfill a distinct role in tracheal stability, as well as respond differently to mechanical stimuli, which collectively gives the trachea its unique anisotropic mechanical properties (Teng et al., 2012). The trachea's horse-shoe shaped hyaline cartilage rings, smooth muscle and annular ligament, collectively produce an anisotropic tissue that allows for longitudinal extensibility and lateral rigidity (Teng et al., 2008; Teng et al., 2012; Boazak and Auguste, 2018). Therefore, any tracheal substitute must be mechanically robust in order to withstand intra-thoracic pressure changes that occur during respiration (Rose et al., 1979; Maziak et al., 1996). Conversely, they must also be able to deform radially to allow for changes in the cross-sectional area during coughing and swallowing (Boazak and Auguste, 2018). These complicated native tissue characteristics, coupled with a lack of standardised protocols to accurately quantify tracheal biomechanics as guidance for implant design, constitute a significant hurdle for tracheal biomaterial scaffold fabrication.

Heterogenous methods to characterise native tracheal mechanics include different specimen types, shapes, testing methods, and parameters suffer from inconsistency [reviewed in Soriano et al. (2021)]. This lack of consistency is likely due to the geometry of the trachea, together with its unique structural regions, which make it difficult to define the most appropriate mechanical tests to perform to best capture the correct biomechanics (Lambert

et al., 1991; Rains et al., 1992; Roberts et al., 1997; Hoffman et al., 2016; Siddiqi, 2017). The inability to collectively define the mechanical properties of the trachea has resulted in studies simply using uniaxial compression testing alone to characterise their scaffolds, which upon in vivo implantation, have subsequently failed due to the mismatch in mechanical properties provoking graft collapse or dislodgement (Grimmer et al., 2004; Wood et al., 2014). To this end, we propose to establish a benchmark of different mechanical test types that are able to evaluate the bulk radial compressive properties of our scaffold but also their flexibility and resistance to applied static load over time, ultimately to recapitulate the mechanical environment of the trachea.

3D-printing (3DP) provides the capacity to generate complex and custom designs that enable the fabrication of patient-specific constructs. Furthermore, due to its reproducibility and the ability to alter mechanical properties of printed constructs through the fine-tuning of complex design features, 3DP was selected in this study to fabricate a polymeric tubular framework for a composite tracheal scaffold. To generate a structure that can produce a tracheal scaffold with mechanical support similar to native cartilage rings, polycaprolactone (PCL) was selected as the printing material due its strong mechanical properties, non-cytotoxic degradation products and long-term stability (Sun et al., 2006; Woodruff and Hutmacher, 2010). These properties have meant that the use of 3DP PCL in tracheal replacement efforts has been widespread, with studies fabricating tracheal scaffolds that display excellent resistance to compressive stresses and facilitate cartilage tissue formation (Gao et al., 2017; Best et al., 2018; Les et al., 2019; Xia et al., 2019; She et al., 2021). However, 3DP PCL constructs in vivo have also suffered from poor results due to inflammatory responses that cause granulation tissue formation and stenosis (Lin et al., 2009; Gao et al., 2017; Rehmani et al., 2017; Kaye et al., 2019; Xia et al., 2019). Notably, in a recent study of a 3D-printed scaffold coated in a chondrocyte suspension and implanted within a rabbit model, all animals had died by 10 weeks, with 75% of deaths occurring linked to excess granulation tissue. Furthermore, no re-epithelisation was found to have occurred (Kaye et al., 2019). Thus, while PCL holds great potential as a base material for mechanically robust tracheal scaffolds, additional components are needed to augment the construct's integration to its surrounding mucosal and submucosal tissue.

The other key issue facing tracheal replacement is that of vascularisation of the implant. This is a critical consideration to avoid tissue necrosis in the implant site and to support new cartilage and epithelium formation (Walles et al., 2004; Li et al., 2019). Tracheal revascularisation—or, more precisely, the lack thereof—was deemed to be a key factor in a high prolife case of scientific misconduct in this field, which involved the implantation of solid polymer tubes seeded with stem cells as tracheal replacement (Delaere and Van Raemdonck, 2016; Delaere et al., 2019). These studies highlighted the importance of the trachea's blood supply, and the need for any substitute graft to effectively support revascularisation and ensure clinical success. Pre-vascularisation prior to implantation has shown higher survival rate of animals, lower rates of stenosis and increased re-epithelisation, suggesting the presence of an established vascular network enhances graft survivability (Luo et al., 2013; Maughan et al., 2017; Wu et al., 2017; Bae et al., 2018).
<figure><img src="image_89.png" /><figcaption>FIGURE 1</figcaption></figure>

Design iterations of 3DP tracheal scaffolds. Two main design features were fabricated: Tubular T; (A) and Partial Ring PR; (B).

<table><caption>TABLE 1 Dimensions and structural differences of tracheal scaffold designs.</caption><thead><tr><td>Scaffold Design</td><td>Dimensions (Inner Ring Radius, mm)</td><td>Dimensions (Outer Ring Radius, mm)</td><td>Opening (degrees)</td><td>Number of Struts</td></tr></thead><tbody><tr><td>Tubular (T)</td><td>4.8</td><td>6.8</td><td>—</td><td>5</td></tr><tr><td>Partial Ring (PR)</td><td>4.8</td><td>6.8</td><td>72</td><td>5</td></tr></tbody></table>

Indeed, our in-house approach to engineer a prevascularised within collagen scaffolds with co-culture of human umbilical vein endothelial cells (HUVECs) and human mesenchymal stem cells (hMSCs) provided *in vitro* and *in vivo* stable vessel formation (Hoffman et al., 2016).

Therefore, we propose that the integration of a 3DP PCL framework within a pre-vascularised collagen matrix can enhance implant integration into surrounding tracheal tissue, as a biocompatible substrate for successful cellular attachment and growth. In this study, building on our bilayered collagen-hyaluronic acid (CHyA-B) matrix (O'Leary et al., 2016), our objective was to incorporate a 3DP PCL framework to produce a reinforced composite CHyA-B scaffold, and demonstrate successful pre-vascularisation as a proof-of-concept prior to future preclinical animal studies.

# 2 Materials and methods

## 2.1 Design and manufacture of 3D-printed PCL framework

Fused deposition modelling 3D-printing was utilised to fabricate novel tubular frameworks with the aim to optimise printing parameters and to mechanically characterise scaffolds for tracheal replacement. The frameworks were designed using custom G-code developed from Python programs, which contain toolpaths for the print head, to generate a design with either a complete 360° tubular ring (T) or 288° partial ring (PR). Both designs featured two rings, 2 mm apart, connected by five struts (Figure 1; Table 1). The designs were 3D printed using an Allevi II Bioprinter (Allevi, United States)

with 25 kDa PCL (Polysciences, Germany). PCL was heated to 90°C and extruded at a pressure of 60 psi at a printing speed of 260 mm/min. All designs were printed using a 25G metal needle (Micron-S, Fisnar, ECT Adhesives, Ireland).

## 2.2 Mechanical testing

### 2.2.1 Radial compression

To determine the mechanical strength of the different designs and the influence of molecular weights of PCL, radial compression was performed using a universal mechanical testing machine Z005 (Zwick/Roell, German, A702416) with a 50 N load cell. The 3DP frameworks were compressed radially at an applied strain of 50% at 0.6 mm/min. The yield point was calculated from the load-deformation curves using the 0.2% offset method. The partial ring designs (PR1-3) were tested in two different testing positions: lateral (L) and anterior-posterior (A-P) to assess the influence of the opening gap on mechanical strength.

### 2.2.2 Cyclical testing

To assess the durability of the different designs, the 3DP frameworks were subjected to cyclical testing. With the universal mechanical testing machine and 50 N load cell, the strain rate was adjusted to mimic a fast-breathing rate of adult humans at 20 breaths per minute for 250 cycles. The designs were placed vertically onto the machine and fine grit sandpaper was placed between the framework and the testing platens to prevent slippage. The peak load at cycles 1 and 250 was measured and the percentage cyclical strain recovery was calculated to elucidate the design's
A Materials:

<figure><img src="image_90.png" /></figure>

Bottom and top
stainless-steel mould
plates

<figure><img src="image_91.png" /></figure>

3D-Printed
Scaffold

<figure><img src="image_92.png" /></figure>

CHyA Film

B
Fabrication Process

<figure><img src="image_93.png" /></figure>

CHyA film wrapped
around peg

<figure><img src="image_94.png" /></figure>

<figure><img src="image_95.png" /></figure>

Top plate secured to
bottom to create a well

<figure><img src="image_96.png" /></figure>

<figure><img src="image_97.png" /></figure>

CHyA slurry
pipetted into well

<figure><img src="image_98.png" /></figure>

<figure><img src="image_99.png" /></figure>

3D-printed scaffold
placed into mould

Freeze Dried

<figure><img src="image_100.png" /></figure>

CHyAB+PCL
trachea scaffold

<caption>FIGURE 2</caption>

Schematic overview of fabrication of reinforced composite CHyA-B tracheal scaffolds. (A) Materials required for scaffold fabrication. (B) Fabrication process of reinforced composite CHyA-B tracheal scaffolds. In brief, CHyA films are wrapped around the pegs of the bottom mould plates and top plates is secured on forming a well in which CHyA slurry is pipetted into. Following this, the 3D-printed frameworks are placed into the mould and freeze dried to form a reinforced composite CHyA-B tubular scaffold.

ability to recovery to its original shape. The percentage of strain recovery after 250 cycles was determined using Eq. 1, where ε<sub>m</sub> is the applied strain and ε<sub>p</sub>(N) is the strain after N cycles (Lendlein et al., 2001).

$$
R_r = \frac{(\varepsilon_m - \varepsilon_p(N))}{\varepsilon_m} \quad (1)
$$

2.2.3 Three-point bend testing

To examine the flexibility of the designs, a 3-point bending rig was used to assess the bending stiffness of the frameworks. In order to provide comparable data, flexural moduli of the scaffolds were assessed as seen in literature for tracheal scaffolds (Park et al., 2018). Frameworks were compressed at 50% strain until failure with the universal mechanical testing machine and 50 N load cell. The partial ring designs were again assessed in two different orientations, firstly with the opening facing down towards the lower plate and secondly with the opening facing out sideways, to assess the influence of the gap on the flexibility of the design.

2.3 Manufacture and characterisation of
reinforced composite CHyA-B scaffolds

2.3.1 Collagen-hyaluronate slurry and film preparation

Collagen-hyaluronic acid (CHyA) slurry and films were prepared using a previously described method (O'Leary et al., 2016). In short, 0.5% w/v collagen (type 1 collagen, bovine tendon) (Collagen Solution, United Kingdom) and 0.044% hyaluronic acid (Hyaluronic acid sodium salt from Streptococcus equi; Sigma-Aldrich, Arklow, Ireland) in 0.5 M

acetic acid were homogenously mixed using an Ultra Turrax T18 Overhead Blender (IKA Works Inc., United States) at 15,000 rpm at 4°C for 3.5 h, and subsequently degassed to 4,000 mTorr under a vacuum to produce a CHyA slurry, which was stored at 4°C. To manufacture CHyA films, 50 mL of the CHyA slurry was cast into 12.5 cm² × 12.5 cm² stainless steel moulds clamped to a polytetrafluoroethylene (PTFE) plate and dehydrated overnight under a constant air flow to produce a thin and transparent CHyA film.

2.3.2 Reinforced composite CHyA-B scaffold manufacture

Prior to incorporating the 3D-printed frameworks to CHyA slurry for freeze-drying, the constructs were surface treated with 3 M sodium hydroxide (NaOH) for 48 h at room temperature, and then rinsed three times with diH20 to remove any residual NaOH. Following the surface treatment, a custom stainless-steel mould was used to fabricate the reinforced composite CHyA-B scaffolds (Figure 2). The mould consists of a bottom plate with 16 pegs, 26 mm in height and 7.8 mm in diameter and a top plate with 14 mm wide circular holes. First, CHyA films were cut into strips of 26 mm × 30 mm and rehydrated in 0.5 M acetic acid for 1 h at room temperature. Upon rehydration, the films were wrapped around the pegs of the bottom plate and left overnight under a constant air flow to dehydrate the films, enabling them to firmly attach to the metal pegs. The following day, the top plate of the mould was then secured on, creating a well in which 1,750 µL of CHyA slurry was pipetted into and then 3D-printed frameworks were then gently placed into the slurry. Subsequently, the films were allowed to rehydrate in the collagen slurry within the mould for 2 h and then freeze-dried using a customised anneal freeze drying cycle. The scaffolds were subjected to an initial freezing step to -20°C for 1 h in which ice crystals formed within the collagen slurry. Following this, the
temperature was raised to -10°C, an annealing step, which merged the ice crystals resulting in larger pores. Finally, to allow for sublimation and drying of the scaffolds, the pressure was decreased to 200 mTorr.

## 2.4 Scaffold ultrastructure

Reinforced composite CHyA-B scaffolds were examined using scanning electron microscopy (SEM) in order to evaluate the collagen integration and the ultrastructure between the PCL fibres. Samples were mounted to an aluminium stub and sputter-coated with gold/palladium (Cressington 108auto, Cressington Scientific Instruments, United Kingdom) at a current of 40 mA for 80 s. Imaging of the scaffolds was performed using a Zeiss Ultra Plus scanning electron microscope (Zeiss, Germany). Images were captured at 5 kV using the secondary electron mode. Three scaffolds from two batches of reinforced CHyA-B scaffolds were used for imaging.

## 2.5 Pore size analysis

Having established the 3DP design and mechanical properties of the scaffolds, the scaffold's ability to support cell infiltration and growth was assessed by measuring the mean pore size of the collagen sub-layer of composite scaffolds. Pore size of the reinforced composite and non-reinforced CHyA-B matrix were analysed to assess the effect of the incorporating tubular 3D-printed framework on the mean pore size of CHyA-B matrices. Samples of reinforced CHyA-B scaffolds were cut into 1 cm × 1 cm sections and analysed using a technique optimised for pore size analysis of collagen scaffolds (O'Brien et al., 2005). In short, the scaffolds were dehydrated using a series of increasing concentrations of ethanol and then embedded in JB-4® glycolmethacrylate (Polysciences Europe, Eppelheim, Germany). Sections of 10 µm thickness of embedded samples were sectioned using a carbon steel blade (C35, PFM, Laboratory Instruments and Supplies, Ireland) on a microtome (Leica RM 2255, Leica, Germany). Sections were stained using 0.5% toluidine blue for 4 min and imaged at ×10 magnification using a microscope (Eclipse 90i, Nikon, Japan), with attached camera (DS R1, Nikon, Japan). Captured images were analysed using a custom-made MATLAB (MathWorks Inc., MA, United States) script developed by our group in conjunction with the Sigmedia Research Group in the Electrical Engineering Department at Trinity College Dublin, Ireland. Images were converted to binary form (black and white) and using a best fit of elliptical lengths the average pore diameter was measured. For each sample, sections were taken every 100 µm spanning the entire length of the scaffold and three scaffolds of each group from two batches of freeze-drying were analysed.

## 2.6 HUVEC and hMSC co-culture on reinforced composite CHyA-B scaffolds

### 2.6.1 Scaffold fabrication for cell culture

In order to biologically characterise the reinforced composite CHyA-B scaffolds, 6 mm × 6 mm sections were 3D-printed as

previously described (Section 2.3) and incorporated and freeze-dried with CHyA slurry and film (Section 2.3.2; Figure 3). Non-reinforced CHyA-B matrices were used as control and were sectioned from a tubular non-reinforced CHyA-B matrices.

### 2.6.2 Sterilisation and crosslinking

All scaffolds for cell study were placed under 254 nm UV light for 30 min on each side for sterilisation at a distance of 10 cm before being chemically cross-linked using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDAC) (Haugh et al., 2011; Lemoine, 2020). Scaffolds were firstly pre-hydrated in Dulbecco's Phosphate Saline (DPBS) before being submerged into a solution of 6 mM EDAC per gram of collagen was mixed with N-hydroxysuccinimide (NHS) at a molar ratio of 2.5 M EDAC: 1 M NHS in 70% ethanol. The scaffolds were then washed three times with sterile DPBS to remove residual cytotoxic product. Scaffolds were stored at 4°C for up to a week after crosslinking or used immediately for in vitro experiments. All steps were performed under sterile conditions.

### 2.6.3 Cell sources and media

All cells present in this study were obtained from Lonza. hMSCs were originally isolated from bone marrow aspirates obtained from the iliac crest of healthy human donors by the commercial supplier with ethical approval (Lonza, Switzerland). HUVECs were commercially acquired (pooled, CC-2719; Lonza, Switzerland). Cells were cultured at 37°C and 5% CO2 under humidified atmosphere. hMSC growth medium corresponded to low glucose (1 g/L) Dulbecco's modified Eagle's medium supplemented (LG-DMEM, Sigma-Aldrich, Ireland) with 10% foetal bovine serum (FBS; Biosera, Ringmer, United Kingdom) and 1% penicillin/streptomycin. HUVEC growth medium corresponded to endothelial growth medium-2 (EGM-2, Lonza, Switzerland). Medium was replaced every 3 days, and upon reaching 80%-90% confluence, cells were passaged using trypsin-EDTA solution. For all experiments described below, the highest passage number used for both cell types were passage 5.

### 2.6.4 Scaffold seeding

To evaluate the influence of 3DP PCL framework on *in vitro* vascularisation of reinforced composite CHyA-B scaffolds, a previously optimised vascularisation cell culture model was utilised, compromising of a co-culture of HUVECs and hMSCs at a ratio of 4:1, respectively (McFadden et al., 2013). All chemically crosslinked and sterilised scaffolds were preconditioned overnight in EGM-2 prior to seeding. Scaffolds were then seeded in a delayed stepwise manner on their porous collagen sub-layer, using 25 µL of cell suspension containing 2 × 10⁵ HUVECs in a dropwise manner. The samples were subsequently incubated for 1 h to allow for cell attachment, followed by the addition of EGM-2. Three days post-HUVEC seeding, scaffolds were seeded with 5 × 10⁴ hMSCs in 25 µL, using the same approach.

## 2.7 Cell viability

To assess and compare cellular viability of the co-culture cells on scaffolds, a battery of tests was performed over a 14-day period using
<figure><img src="image_101.png" /><figcaption>FIGURE 3</figcaption></figure>

Scaffold design for cell culture study, the g-code script was altered to reduce the circumference and height of the scaffold to retain the (A) strut as (B) a stacked 6 × 6 mm scaffold.

assays. An alamarBlue™ metabolic assay (Invitrogen, BioSciences, Ireland) and PicoGreen dsDNA assay (Invitrogen) were performed at days 6, 10 and 14 of culture, in addition to fluorescent imaging of live and dead cells at day 14 using LIVE/DEAD® Viability/ Cytotoxicity Kit (Thermofisher, Ireland) all according to the manufacture's protocol.

## 2.8 Vessel formation in reinforced composite CHyA-B scaffolds

### 2.8.1 Immunofluorescence

To image the formation of vessel-like structures within scaffolds, scaffolds were harvested at days 6, 10 and 14, for fluorescence labelling (do Amaral et al., 2019a). The scaffolds were washed with PBS and fixed in 10% neutral buffered formalin solution (Sigma-Aldrich, Ireland) for 1 h at room temperature. Following fixation, scaffolds were washed three times with PBS and cut in half transversely to expose the cross-section for imaging. 0.3% Triton X-100 in PBS followed by 3% bovine serum albumin (BSA; Sigma-Aldrich, Ireland) in PBS were applied for 20 min each to permeabilise cells and block unspecific antigen binding. Monoclonal mouse anti-human CD31 antibody (Dako, M0823) diluted to 1:50 in 1% BSA in PBS was used for HUVEC identification and a rabbit polyclonal antibody for α-smooth muscle actin (α-SMA; Abcam, ab5694) diluted at 1:100 in 1% BSA in PBS was applied to detect hMSCs. Primary antibodies were incubated overnight at 4°C. Rat anti-mouse IgH (H + L) secondary antibody, Fluorescein isothiocyanate (FITC; Biosciences, Ireland, 11401185), diluted at 1:100 in 1% BSA in PBS and goat anti-rabbit IgG (H + L) cross-adsorbed secondary antibody, Alexa Flour 633 (Invitrogen, A-21070), diluted at 1:400 in

1% BSA in PBS was applied for 1 h at room temperature. Thereafter, 1:600 diluted Phalloidin-Tetramethylrhodamine B isothiocyanate (TRITC; Sigma-Aldrich, Ireland) in 1% BSA in PBS was applied for 20 min to stain the cell cytoskeleton followed by 4', 6-Diamidino-2-phenylindole dihydrochloride (DAPI; Sigma-Aldrich, Ireland) at 1: 1,000 dilution in PBS for 20 min for nuclei staining. Samples were stored in PBS at 4°C until imaging with a Carl Zeiss LSM 710 confocal microscope with a W N-Achroplan 10x (numerical aperture 0.3) lens. Z-stack images were taken at 30 µm depths below the surface to yield a total 150 µm total depth.

### 2.8.2 Chick chorioallantoic membrane (CAM) model

To assess the influence of the 3DP PCL framework within the CHyA-B matrix on vessel infiltration *ex vivo*, scaffolds were incubated within a chick chorioallantoic membrane (CAM) model as previously described (do Amaral et al., 2019a; Ryan et al., 2019; do Amaral et al., 2019b). Fertilised chicken eggs (Ovagen Group Ltd., Co. Mayo, Ireland) were purchased at day 0 of development and incubated at 37°C for 3 days. Following incubation, the eggs were cracked into 100 mm diameter petri dishes (Corning Inc., New York, United States), which in turn were then placed within larger 150 mm diameter petri dishes containing sterile PBS, to form a humidified chamber. The chicks were incubated for a further 4 days, completing 7 days of total development, viability of embryos was firstly ensured and checked at every day from this point forward; when cross-linked and sterilised CHyA-B matrices and reinforced composite CHyA-B scaffolds were placed within the chick membrane and incubated for a further 5 days, allowing for vessel ingrowth. At day 12 the samples were harvested and imaged. The vascularisation around the scaffolds was quantified using ImageJ software after treatment with the “Mexican Hat Filter” to outline
the blood vessels, and then converted to 8-bit and the blood vessel area was measured.

## 2.8.3 Histology

Harvested scaffolds from CAM were stained with haematoxylin and eosin (H&E) to observe vessel infiltration throughout the scaffold. Scaffolds were fixed overnight in 10% neutral buffered formalin solution at 4°C. Following fixation, scaffolds were dehydrated in increasing concentrations of ethanol until 100% ethanol was reached. Scaffolds were then placed within xylene solution under agitation, to dissolve out PCL filaments, after which they were processed using an automated tissue processor (ASP300, Leica, Germany) for paraffin infiltration. The samples were then embedded into paraffin blocks and were sectioned as described in Section 2.5 to obtain sections with a thickness of 7 µm in their cross-section to observe vessels infiltration. Sections were rehydrated in decreasing concentrations of ethanol, and then incubated with Harris haematoxylin (Sigma-Aldrich, Ireland) for 5 min before being washed in tap water for 5 min. Differentiation of the samples was then performed with acidified 70% ethanol, before being stained with 0.1% eosin Y (Sigma-Aldrich, Ireland) in 95% ethanol. Finally, slides were dehydrated and mounted with DPX. Images were captured as described in Section 2.5.

## 2.8.4 Pro-angiogenic protein expression

The expression of pro-angiogenic markers vascular endothelial growth factor (VEGF) and basic fibroblast growth factor (b-FGF) were quantified using ELISA kits (R&D, Biotechnhe, Ireland). Supernatants on days 6, 10 and 14 were collected and analysed following the manufacturer's protocol.

## 2.8.5 Pro-angiogenic gene expression

To assess gene expression within scaffolds seeded with the co-culture model, scaffolds were harvested and stored at -80°C in 1 mL of Qiazol lysis reagent (Qiagen, Crawley, United Kingdom). Total RNA was isolated using a RNeasy Mini Kit according to manufactures protocol (Qiagen). RNA quality and quantity was determined using Nanodrop 2000 Spectrophotometer (Thermofisher, Ireland). Reverse transcription of RNA lysates was performed on 400 ng of total RNA using the QuantiTect Reverse Transcription kit (Qiagen) according to manufacturer's protocol on a MiniAmp Thermal Cycler (A37834, Thermofisher). Quantitative real time polymerase chain reactions were performed in duplicate on a 7500 Real-Time PCR system (Applied Biosystems) using QuantiTect SYBR Green PCR kit (Qiagen). The mRNA relative expression was calculated by 2⁻ΔΔCT method, in which CHyA-B matrices were used as control relative to each respective timepoint. Target mRNAs analysed were KDR (QT00069818, Qiagen), ANG1 (QT00046865, Qiagen), ANG2 (QT00100947, Qiagen), TEK2 (QT01666322, Qiagen), FLT1 (QT00270823, Qiagen) and 18S (QT00199367, Qiagen) was used as a housekeeping gene.

## 2.9 Data analysis

Quantitative data were analysed using GraphPad Prism 8.0 software. Two-tailed student t-tests were used for statistical analysis between two groups, and 1-way or 2-way ANOVA was

used for statistical analysis between multiple groups. Sidak's multiple comparison post hoc analysis was performed in all ANOVA assessment.

# 3 Results

## 3.1 Scaffold characterisation

### 3.1.1 Mechanical testing

To determine the mechanical strength of the two designs, uniaxial compressive testing was carried out (Figure 4). A significant increase in the yield load was calculated in the tubular design, 5.177 ± 0.219 N/mm, and the A-P and L partial-ring designs, 5.523 ± 0.744 N/mm and 1.313 ± 0.061 N/mm respectively, compared to CHyA-B matrices, 0.307 ± 0.453 N/mm. Notably, a difference in yield load was observed between the two orientations of the PR design, suggesting this design better recapitulated the anisotropic mechanical properties found in native tracheal tissue.

Moreover, to assess the durability of the different designs, and whether the 3D printed scaffolds could withstand the constant mechanical loading and unloading that the trachea experiences during respiration, all scaffold designs underwent cyclical testing. Irrespective of design feature, all scaffold designs withstood cyclical loading over 250 cycles and remained patent, with no visible cracks or failure sites observed when inspected visually. However, a significant 2N difference in average peak load was observed between tubular and partial ring design at cycle 1, which reduced to a 1N difference at cycle 250, highlighting the influence of the partial ring design when under radial mechanical stresses (Figure 4B). Nonetheless, the difference in structural design did not influence the percentage cyclical strain recovery (Figure 4C). Scaffolds recovered an average of 67% of the applied strain following 250 cycles to 15% applied strain in the tubular design and 71% in the partial ring design, significantly greater than the non-reinforced CHyA-B matrix.

To determine the flexibility of the scaffolds, both designs underwent three-point bend testing. The partial ring design was tested under two different orientations in relation to the position of the opening gap (Figure 4D). The tubular design exhibited a higher flexural modulus of 0.153 ± 0.010 MPa. Additionally, the partial ring design also exhibited a similar high flexural modulus in the lateral position (0.131 ± 0.005 MPa) but was significantly more flexible in the second testing position A-P (0.266 ± 0.022 MPa).

### 3.1.2 Scaffold ultrastructure

Having successfully mechanically characterised the 3DP designs, we next sought to evaluate the ultrastructure of the reinforced composite CHyA-B scaffolds after incorporating the 3DP frameworks into CHyA-B matrices using SEM. Scaffolds were imaged using SEM to assess the influence of the PCL fibres and scaffold design on collagen matrix infiltration. Subsequent SEM assessment captured successful integration of collagen within the PCL fibres (Figure 5). Therefore, a favourable structure was achieved through a reproducible process achieving successful integration of a porous collagen network around a 3D-printed backbone structure.

### 3.1.3 Pore size analysis

To assess the scaffold's ability to support cellular infiltration and growth, the mean pore size of the collagen matrix in the sub-layer of
<figure><img src="image_102.png" /></figure>

<figure><img src="image_103.png" /></figure>

<figure><img src="image_104.png" /></figure>

<figure><img src="image_105.png" /><figcaption>FIGURE 4</figcaption></figure>

Mechanical properties of 3DP tracheal framework designs. Scaffolds were subjected to multimodal mechanical assessment to compare the influence of different scaffold designs. (A) Compression testing, (B, C) cyclical testing, (D) Three-point bending. Results displayed as mean ± SEM, *n* = 4.
ns = *p* > 0.05, ***p* < 0.001, ****p* < 0.0001.

<figure><img src="image_106.png" /></figure>

<figure><img src="image_107.png" /></figure>

<figure><img src="image_108.png" /></figure>

<figure><img src="image_109.png" /><figcaption>FIGURE 5</figcaption></figure>

(A–C) Scanning electron micrographs showing the ultrastructure of cross-sectioned reinforced composite CHyA-B scaffolds. PCL indicated with red arrow. (D) A representative surface image of the non-film outer collagen surface layer showing the porosity (blue arrow) of the collagen sub-layer. *n* = 2, in triplicate.
<figure><img src="image_110.png" /><figcaption>FIGURE 6</figcaption></figure>

Pore size analysis of bilayered composite scaffold. (A) Mean pore diameter of non-reinforced (CHyA-B) matrices and reinforced composite (CHyA-B+PCL) scaffold. (B) Temperature profile of tubular moulds during –20 anneal cycle in comparison to shelf temperature, data captured by Dr. Derek Whelan, RCSI (n = 1). Toluidine blue sections of CHyA-B (C) and CHyA-B+PCL (D) scaffolds used for software analysis Scale bar = 100 µm. Results displayed as mean ± SEM, n = 3, in quadruple. ****p < 0.0001.

reinforced composite CHyA-B scaffolds was analysed. The inclusion of the 3D-printed PCL framework within CHyA-B tubular matrix resulted in an average pore size of 290 ± 27.66 µm compared to an average pore size of 180 ± 1.962 µm in non-reinforced tubular CHyA-B matrices (Figure 6A). Temperature probes placed within the tubular moulds and freeze-dryer shelf recorded differences in freeze drying profiles (Figure 6B), with the temperature within the tubular moulds requiring a longer time to be brought up back to temperature during the cycle, possibly influencing the pore size of the resultant scaffolds (Figures 6C, D).

## 3.2 Cell viability

In an aim to assess and compare cell viability of the vascularisation co-culture system on reinforced CHyA-B scaffold to that of CHyA-B matrix, scaffolds were subjected to multiple viability assays (Figure 7). Cell metabolic activity measured in reinforced composite CHyA-B scaffolds was comparable to measurements in control CHyA-B matrices over 14 days of culture period (Figure 7A). Notably, double stranded DNA (dsDNA) content was two-fold greater in reinforced composite CHyA-B scaffolds compared to CHyA-B matrices over 14 days of cell culture (Figure 7B). LIVE/DEAD® imaging also confirmed the presence of a high population of green viable cells on both scaffold types (Figure 7C). Overall, the mix of quantitative and qualitative data indicate that the inclusion of 3DP PCL framework

within CHyA-B matrices sustains vascular cell viability, and potentially facilitates increased proliferation.

## 3.3 Vessel formation in reinforced composite CHyA-B scaffolds

### 3.3.1 Immunofluorescence

To image the formation of vessel-like structures within scaffolds, cells were fluorescently labelled and then imaged using confocal microscopy. The presence of vessel-like structures was visible in reinforced composite CHyA-B scaffolds and CHyA-B matrices across 14 days of culture (Figure 8; white arrows), with fewer vessel-like structures on the post-peak timepoint of day 14. Furthermore, fluorescence labelling of endothelial cell makers CD31 and pericyte cell marker α-SMA on day 10 of cultured samples (Figure 9) confirmed the presence of both cell types around vessel-like tubules. Thus, we confirmed that the incorporation of 3DP PCL framework within CHyA-B matrix did not impede of vessel formation in vitro.

### 3.3.2 Pro-angiogenic protein expression

The expression of pro-angiogenic markers VEGF and b-FGF was quantified using ELISA kits on days 6, 10 and 14 of in vitro culture. Although no significant difference was observed between reinforced composite CHyA-B scaffolds and CHyA-B matrices; over
<figure><img src="image_111.png" /><figcaption>FIGURE 7</figcaption></figure>

Assessment of cell viability of reinforced composite CHyA-B scaffolds. CHyA-B matrices and reinforced composite CHyA-B scaffolds were subjected to cell viability tests to assess their biocompatibility with HUVECs and hMSCs. (A, B) Cellular metabolic activity and double stranded DNA content measured in reinforced composite scaffolds (CHyA-B+PCL) and non-reinforced matrix (CHyA-B). Representative LIVE/DEAD® images taken of the surface of CHyA-B (C) and CHyA-B+PCL (D) scaffolds at day 14 at ×4 magnification. Scale bar = 100 µm. Results displayed as mean ± SEM, *n* = 3, in triplicate. ns = *p* > 0.05, ***p* ≤ 0.001.

the 14 days of culture period, comparable levels of VEGF and b-FGF was measured (Figure 10).

### 3.3.3 Pro-angiogenic gene expression

The influence of reinforced composite CHyA-B scaffolds on HUVEC and hMSC co-culture gene expression of KDR, TEK 2, ANG 1, ANG 2, and FLT 1 was analysed to assess the effect of the 3DP PCL framework on angiogenic gene expression. The inclusion of the 3DP PCL framework was found to not significantly alter the expression of angiogenic genes (Figure 11), with gene expression sustained over 14 days of *in vitro* culture. Taken together with the expression of pro-angiogenic proteins (Figure 10), this work demonstrates the mechanically-reinforced tracheal scaffold's ability to sustain angiogenic marker expression *in vitro*.

### 3.3.4 Chick chorioallantoic membrane (CAM) model

To assess the influence of the 3DP PCL framework on vessel infiltration *ex vivo*, scaffolds were analysed using a CAM model, in which they were placed within fertilised chick eggs and incubated to allow for living blood vessels to infiltrate. Image analysis of macro-images of harvested samples (Figure 12A) calculated an increase in vascularised area in the reinforced composite CHyA-B scaffolds over its non-reinforced counterpart, although not statistically significant

(Figure 12B). Furthermore, sectioned samples of reinforced composite scaffolds also confirmed the presence of vessels penetrating the scaffolds confirming that the incorporation of PCL fibres did not disrupt vessel infiltration (Figures 12C–F). These are important findings as we have shown the successful fabrication of a novel mechanically reinforced tubular scaffold for tracheal tissue regeneration, with the ability to successfully vascularise *in vitro* and *in vivo*, which would enhance graft survivability post implantation.

## 4 Discussion

In this study, we sought to fabricate a composite biomaterial-based scaffold for the regeneration of damaged tissue in large tracheal defects. The major objective of the study was to investigate the potential of fabricating reinforced CHyA-B scaffolds tubular constructs with similar mechanical properties to native tracheal tissue by incorporating a 3D-printed PCL framework within the porous collagen sub-layer. In particular, we assessed the feasibility of 3D-printing tubular constructs, evaluated the mechanical properties of different design, and validated their *in vitro* biocompatibility and then the ability to successfully vascularise. Multi-modal mechanical assessment led to the
<figure><img src="image_112.png" /><figcaption>FIGURE 8</figcaption></figure>

Actin cytoskeleton and nuclei staining on reinforced composite CHyA-B scaffolds and CHyA-B matrices *in vitro* vascularisation. HUVECs were cultured with hMSCs on either (A–C) non-reinforced matrices (CHyA-B) or (D–E) reinforced composite scaffolds (CHyA-B+PCL) scaffolds for 14 days. Maximum intensity projections reconstructed from z-stack images display cells arranged in tubule-like structures (white arrows). Magenta = f-actin, blue = nuclei. *n* = 3, in triplicate. Scale bar = 100 µm.

mechanical characterisation of the 3DP framework to achieve two
lead designs (tubular and partial ring) with mechanical properties
comparable to native tracheal tissue. Furthermore, the incorporation
of the 3DP framework and tubular form-factor generated a scaffold
with greater mean pore diameter, ideal for cellular infiltration.
Biological characterisation of the reinforced composite CHyA-B
scaffolds demonstrated their ability to support cellular growth.
Additionally, vascularisation within the PCL-reinforced collagen
matrix was maintained with no adverse outcome; vessel-like
tubules were observed repeatability in scaffolds using a suite of
methods. The expression of pro-angiogenic markers was also
investigated, to examine the effect of 3DP PCL on protein and
gene expression, with no significant difference observed to the
standard vascularised scaffold controls. Finally, in the first step to
*in vivo* validation of the scaffolds, the CAM model was utilised and

confirmed sustained levels of vascularisation within the reinforced
composite scaffolds compared to CHyA-B matrices. Collectively, the
data presented within this chapter confirmed that the incorporation
of 3DP PCL for mechanically reinforcing CHyA-B matrices
maintained successful vascularisation within the collagen matrix
and represented a step towards a functional, multi-cellular, and
viable construct for tracheal regeneration. Taken together, the
results demonstrate that a composite CHyA-B scaffold can be
developed to recapitulate the mechanical properties of native
tracheal tissue with no negative effects on cell proliferation and
vascularisation.

Tracheal scaffold designs were conceptualised with the aim to
match the aforementioned mechanical properties of native trachea.
The use of 3DP technology has become more widespread within the
field of bioengineering for customised tissue replacement or
<figure><img src="image_113.png" /><figcaption>FIGURE 9</figcaption></figure>

CD31 and α-SMA expression by HUVECs and hMSCs in co-culture on reinforced composite CHyA-B scaffolds and CHyA-B matrices at day 10 on *in vitro* culture. HUVECs were cultured with hMSCs on either (A–D) non-reinforced matrices (CHyA-B) or (E–H) reinforced composite scaffolds (CHyA-B+PCL). Maximum intensity projections reconstructed from z-stacks display tubule-like structures lined with HUVECs CD31 stained in green; (B, F) and hMSCs in proximity α-SMA stained in red; (C, G). *n* = 3, in triplicate. Scale bar = 100 µm. Nuclei = blue, f-actin = magenta, α-SMA = red, CD31 = green.

regeneration strategies for an array of organ systems. PCL was selected as the printing material due to its excellent printing characteristics. The thermoplastic nature of PCL and its low melting point temperature (60°C–70°C), makes it available for use for a wide variety of 3D printing methods. Furthermore, the mechanical properties of PCL has seen its widespread use in cartilage replacement in tissue engineering, and investigative studies in tracheal tissue engineering (Pfeiffer et al., 2008; Kundu et al., 2013; Hollister et al., 2016; Morrison et al., 2017; Les et al., 2019). Forgoing the solid polymer tubular design commonly used in recent tracheal scaffold fabrication attempts (Pfeiffer et al., 2008; Zych et al., 2012; Gao et al., 2017; O'Sullivan et al., 2019), two rings of PCL were fabricated into either complete 360° tubular ring or a 288° partial ring. The use of struts was hypothesised to not only

generate pores for the collagen matrix to infiltrate within the 3DP framework, but also to provide a polymeric backbone with a degree of flexibility by generating space between the two rings. Also, layer-by-layer stacking of struts and rings were hypothesised to generate mechanical integrity and robustness—thus, producing framework designs that could potentially generate the necessary mechanical strength but also be flexible (Kaye et al., 2019).

To assess whether the 3D-printed tracheal framework designs could match the trachea's anisotropic mechanical properties, all designs firstly underwent radial compressive testing (Figure 4A). Radially compressing samples is often performed on pipes for assessment of resistance to external compressive forces in cases where classical dog bone shaped samples are not feasible, as with curved tubular structures. Unconfined, uniaxial compression testing
<figure><img src="image_114.png" /></figure>

<figure><img src="image_115.png" /><figcaption>**FIGURE 10**
Expression of pro-angiogenic proteins (A) VEGF and (B) b-FGF were measured over 14-day culture period on scaffolds in co-culture of HUVECs and hMSCs on reinforced composite scaffolds (CHyA-B+PCL) and non-reinforced CHyA-B matrices. Results displayed as mean ± SEM. *n* = 3, in triplicate. ns = *p* > 0.05.</figcaption></figure>

was utilised to assess the substrate stiffness of the 3D-printed frameworks radially between the different design features and the molecular weights of PCL. The resistance to compressive forces within the framework designs were greater than reported figures of human trachea which ranged from 1.06 to 1.08 N A-P and 0.55-2.32 N L (Huang et al., 2021), however, a study investigating the difference in mechanical properties within different large animal models and the influence on specimen length on mechanical properties reported figures within range of calculated values of our scaffolds, ranging from 2 to 5 N (Gao et al., 2017; Huang et al., 2021) in complete segments of tracheas tested at 4-8 rings in length at 50% deformation. For tracheal specimens radially compressed in both anterior-posterior and lateral testing positions, values of 2.38 N in canine and 2 N (A-P) and 2 N (L) both at 50% radial compression and 5 N at 40% in a goat trachea have been reported (Gao et al., 2017; Morrison et al., 2017). As already highlighted, the importance of matching the mechanical properties of the native tissue and implanted scaffold is vital, therefore while the mechanical strength of scaffold designs may be too high for human trachea, these designs could still be utilised for animal models. This also further supports the use of 3DP as it would allow for the adjustment of the designs to match mechanical properties of animal model with relative ease. Additionally, the resistance of mechanical forces of our scaffolds is much lower than other tissue-engineering attempts. 3DP PCL scaffolds fabricated for tracheal replacement reported values of approximately 23 N (She et al., 2021) and 30 N (Xia et al., 2019) of full length scaffolds, whereas electrospun tracheal scaffolds have generated values of 32.9 N (L) and 104 N (A-P) (Best et al., 2018) when displaced to 75% of luminal diameter, which are too mechanically stiff compared to native tissue. The 3DP frameworks generated in this study could therefore be beneficial for tissue integration as a function of their reduced mechanical resistance.

Although the trachea has excellent stability and radial strength in order to prevent collapse, it also needs to deform and quickly recover during breathing and coughing (Zhao et al., 2019). In an attempt to investigate similar durability of the PCL frameworks and

also the influence of strut numbers on durability, cyclical loading and unloading was carried out for 250 cycles. Testing the designs in this manner would in theory increase the possibility of delamination occurring between the 3DP layers, leading to a loss of structural stiffness and ultimately failure of the scaffold design (Baumann and Hausmann, 2021). Positively, we found that all framework designs survived 250 cycles at 15% applied strain (Figures 4B, C). Cyclical testing of tracheal scaffolds has not been widely investigated, with researchers usually conducting simple lateral compressive or tensile tests only, however an ovine model tested under cyclical loading over 250 cycles under 15% strain reported a peak load of 0.06 N at cycle 1 and 0.04 N at cycle 250 (Mansfield et al., 2016). Through testing the different designs of our scaffolds, we were able to confirm that the 3DP PCL could withstand radial deformation over a large number of cycles without failure; furthermore, no visible failure sites or delamination were observed. A notable difference was observed between the two main design features, with the tubular design resistant to much higher radial compressive forces than the partial ring design. From the same testing regime of cyclical testing, the cyclical strain recovery was calculated to assess the percentage recovery of all the designs over 250 cycles of compressive loading. Notably, although a large difference in peak load was observed between tubular and partial ring designs, there was no significant difference in percentage recovery between the two main design features. This may be due to the high elastic modulus of PCL itself, which has been reported to correlate to the higher degree of crystallinity of 3DP PCL filaments and resultant higher elastic modulus and percentage recovery (Soufivand et al., 2020).

The final mechanical assessment assessed the flexibility of the 3DP designs by a three-point bending test. To reiterate, the trachea not only needs to radially deform to allow for narrowing and widening for passing of food through the oesophagus and coughing, but also needs to have a high degree of flexibility to allow for movement of the neck with ease (Rains et al., 1992; She et al., 2021). Furthermore, a highly flexible scaffold is less likely to dislodge when a patient suffers from a coughing fit, which has been an outcome in vivo and clinical trials of previous tracheal
A

KDR

<figure><img src="image_116.png" /></figure>

C

ANG 1

<figure><img src="image_117.png" /></figure>

E

<figure><img src="image_118.png" /></figure>

B

TEK 2

<figure><img src="image_119.png" /></figure>

D

ANG 2

<figure><img src="image_120.png" /></figure>

FLT-1

**FIGURE 11**
Expression of pro-angiogenic genes expressed in co-culture of HUVECs and hMSCs on reinforced composite CHyA-B+PCL scaffolds over 14-day culture period. Quantification of mRNA expression of (A) KDR expression (B) TEK 2 expression (C) ANG 1 expression (D) ANG 2 and FLT-1 expression (E) displayed as mean ± SEM with expression relative to mRNA expression in CHyA-B matrices at each relative time point. *n* = 3; in triplicate. *ns* = *p* > 0.05.

replacement efforts and a huge cause for failure (Grimmer et al., 2004; Azorin et al., 2006; Luo et al., 2013; Wood et al., 2014; Elliott et al., 2017). Designs T and PR generated flexural moduli of 0.13–0.26 MPa (Figure 4D), an order of magnitude lower than other tissue-engineering tracheal scaffold studies using 3DP PCL with a flexural modulus of 1.088 MPa (Park et al., 2018), thus making this study’s scaffold design much more flexible and compliant. However, the flexural moduli of rabbit trachea has been reported to range between 0.214 and 0.25 MPa (Pfeiffer et al., 2008; Best et al., 2018). Although no reported literature exists that has analysed the flexural moduli of human trachea, which is likely to be greater than smaller rabbit trachea, a study into the changes in mechanical properties of cell-seeded implanted tracheal scaffolds did find a 60% increase in flexural moduli 8 weeks post implantation from 1.71 to 2.84 MPa (Lin et al., 2008). Compressive and cyclical testing confirmed both second generation framework designs are still able to withstand radial compressive forces and are durable, therefore, higher flexural

moduli could potentially be achieved in either design through matrix remodelling of cell-seeded scaffolds. Next steps would seek further surgeon evaluation of our designs, especially blinded evaluation, which would provide vital feedback on the scaffolds handling for future reiterations of the scaffold.

After establishing final 3DP parameters, designs, and mechanical properties for the tracheal scaffolds, we sought to investigate the biological properties of the composite scaffolds. The 3DP tubular PCL framework was successfully incorporated with CHyA films and slurry to fabricate composite reinforced CHyA-B scaffolds with ease and reproducibility, generating scaffolds with uniform matrix incorporation and tubular form factor. The mean pore size of the porous collagen sub-layer was analysed to investigate whether the incorporation of 3DP PCL within CHyA-B matrix had an effect, if any, on the pore size. The results indicated that the change in form factor of tubular CHyA-B scaffolds from flat 2 mm CHyA-B matrix used as *in vitro* respiratory models increased from 80 µm (O’Leary et al., 2016) to
A

CHyA-B

<figure><img src="image_121.png" /></figure>

CHyA-B+PCL

<figure><img src="image_122.png" /></figure>

B

<figure><img src="image_123.png" /></figure>

C

<figure><img src="image_124.png" /></figure>

<figure><img src="image_125.png" /></figure>

<figure><img src="image_126.png" /></figure>

<figure><img src="image_127.png" /><figcaption>**FIGURE 12**
Reinforced composite CHyA-B scaffolds were vascularised ex vivo in a chick chorioallantoic membrane (CAM) model. (A) Macro-images taken of harvested scaffolds after incubation within CAM. (B) Image analysis of macro-images of harvested scaffolds calculated greater vascularised area within reinforced composite scaffolds (CHyA-B+PCL) than the non-reinforced CHyA-B matrices. (C–F) H&E stained sections of CHyA-B matrices (C, D) and reinforced composite CHyA-B+PCL scaffolds (E, F) imaged at ×4 magnification. Staining confirmed the presence of blood vessels infiltrating throughout the scaffolds as highlighted by blue arrows. Results displayed as mean ± SEM, *n* = 5. ns = *p* > 0.05. Scale bar = 100 µm.</figcaption></figure>

180 µm. This increase in pore size may have been a result of the increase in mass of metal and overall different mould form factor of the tubular moulds in comparison to the smaller flat square moulds. This in turn influenced the freezing temperature within the mould, altering the freeze-drying profile of the –20 anneal cycle of the tubular scaffolds which impacted the pore size (Figure 6B). Moreover, the incorporation of the 3DP PCL framework into the CHyA-B matrix significantly increased the mean pore size from 180 µm to 290 µm (Figure 6A). Literature shows the ideal pore size for vascularisation is around 300–400 µm (Bai et al., 2010), therefore, the incorporation of 3DP PCL within the collagen matrix produced a composite scaffold with a sub-layer of collagen with a pore size ideal for cellular infiltration (Murphy et al., 2010) and vessel formation.

Having confirmed the feasibility of incorporating the 3DP PCL framework within the CHyA-B matrix and mechanically evaluated different designs for a tracheal replacement scaffold, we now sought to investigate the influence of incorporating the 3DP PCL framework on effective vascularisation *in vitro* and in the CAM model. Although collagen-glycosaminoglycan scaffolds are validated for *in vitro* and *in vivo* pre-vascularisation (McFadden et al., 2013), we wanted to ensure that the inclusion of the PCL reinforcement allowed for sustained vascularisation within the collagen matrix. The introduction of 3DP polymer fibres could impede cell-cell interactions, thereby decreasing vessel formation, and was a necessary risk to investigate. As such, the key objective of this study was to confirm the scaffolds' ability to support cell growth of the *in vitro* co-culture system of HUVECs and hMSCs.

As with most 3D cultures, it was important to initially confirm cell viability in the PCL-reinforced scaffolds prior to any extensive analysis of angiogenesis or vascularisation. The composite scaffold exhibited

comparable levels of cellular metabolic activity over 14 days of culture with no significant difference observed (Figure 7A), whereas levels of dsDNA present within the reinforced composite scaffolds were observed to be 2-folds greater than their non-reinforced counterpart (Figure 7B). The incorporation of the 3DP framework to CHyA-B scaffolds increased the mean pore size of the collagen sub-layer from 180 µm to 290 µm (Figure 6), and plausibly facilitated more cellular growth through augmented cell infiltration, nutrient inflow, and waste outflow (Murphy et al., 2010). Differences between the two viability assays may have been due to the differences in application of each assay. With the measurement of dsDNA content, the scaffold is broken down and cells are lysed whereas with measuring the cellular metabolic activity there could be limitations in how far the solution penetrates the scaffolds and thus provides limited measurements of the cells on the periphery of the scaffold. Finally, the presence of a high coverage of green viable cells on the surface of reinforced composite CHyA-B scaffolds was observed when stained with LIVE/DEAD®. Taken together, this further confirmed the biocompatibility of the scaffolds, with sustained cell growth and proliferation maintained with the co-culture model on reinforced composite CHyA-B scaffolds.

It is important to note that the composition of the polymer framework can also have a significant impact on viability, in addition to its effects on composite scaffold porosity. As previously discussed, PCL does not release cytotoxic degradation products into the cell culture environment, and therefore its biodegradation process does not impact cell growth (Sun et al., 2006; Brittberg, 2008). In contrast, other 3DP materials investigated for tracheal tissue replacement such as PLA (polylactic acid) and PGLA (poly(lactic-co-glycolic acid)) have biodegradation products that have been found to trigger an inflammatory response leading to a low pH-environment (Liu and Cao, 2007), and therefore causing an accelerated loss of
structural integrity. For example, this was reported in a 3DP PLA tracheal scaffold which when implanted into the fascia of rabbits for pre-vascularisation, shrunk and deformed due to its instability whereas its PCL counterpart maintained its shape (Tsao et al., 2014). Thus, the addition of the 3DP PCL framework has shown to maintain cell viability while generating mechanical integrity.

The formation of vessel-like tubules was confirmed within reinforced composite CHyA-B scaffolds by several methods. The inclusion of the PCL framework to CHyA-B matrix has been shown to increase the pore size of the collagen matrix (Figure 6), thus it was imperative to assess if the change in pore size could impact vessel formation within the scaffolds. Firstly, confocal microscopy of fluorescently labelled cells within the scaffolds demonstrated the presence of vessel-like tubules for up to 14 days of culture (Figure 8). Notably, we also confirmed that the tubule-like structures were characterised by the lining of HUVECs (CD31) with hMSCs (α-SMA) in direct contact by staining with cell specific markers (Figure 9). As with the previously established model (McFadden et al., 2013), a similar trend was observed, with the strongest vessel formation occurring on day 10 of culture followed by noticeable vessel regression at day 14. As is the issue often associated with long term static in vitro culture of microvascular networks, disengagement of cells can take place, resulting in structural regression of the networks, thus limiting the extent to which vessel development within a scaffold can take place in vitro. However, once implanted in a dynamic environment in vivo, recruitment of host perivascular cells can promote further stabilisation of blood vessels (Montaño et al., 2010).

This possibility was bolstered by the analysis of angiogenic marker expression in cell-seeded scaffolds. No statistical difference seen between the CHyA-B matrix and the reinforced composite CHyA-B scaffolds with regards to expression of angiogenic genes (Figure 11) and proteins (Figure 10), indicating parity in vascularisation potential. The protein bFGF (also known as FGF-2) and its counterpart FGF-1, are potent pro-angiogenic markers which promote the proliferation and differentiation of endothelial cells (Lee et al., 2000; Oladipupo et al., 2014). The expression of both VEGF and bFGF is dynamic during vascularisation (Bautch, 2012; Brundo et al., 2013). The stabilising of KDR (also known as VEGF receptor 2) and FLT-1 (also known as VEGF receptor 1) expression at day 14 but the increase in expression of ANG-1 and its receptor TEK 2 could suggest the maturation of vessel-like structures occurring within the scaffolds from early vessel sprouts to more mature vessel tubules. KDR and FLT-1 are major signal transducer for angiogenesis, promoting vessel formation (Helfrich and Schadendorf, 2011) and when the pro-angiogenic VEGF binds and activates signalling via its receptor KDR on endothelial cells, extracellular matrix degradation, tip cell migration and endothelial proliferation are promoted to form new immature vessel sprouts. Whereas ANG-1 is a stimulator of vessel growth and along with its receptor TEK 2 promotes vessel stabilisation (Lambert et al., 1991; Roberts et al., 1997). However, the upregulation of ANG-2, a marker that causes disruption of vascularisation and destabilisation (Helfrich and Schadendorf, 2011) would suggest the co-culture model is not able to hold up angiogenesis at day 14 as confirmed previously (McFadden et al., 2013) and as shown in fluorescently labelled cultured scaffolds on day 14 imaged samples (Figure 8).

In the first step of *in vivo* validation of the reinforced scaffolds, their vascularisation potential was further validated using a CAM model, with an increased area of vascularisation measured within the reinforced composite CHyA-B scaffolds (Figure 12; p = 0.1929). As with the cellular proliferation, the increased pore size of the reinforced composite CHyA-B scaffolds may have allowed for greater blood vessel infiltration, as a mean pore diameter of 300–400 µm has been found to be ideal for vascularisation (Bai et al., 2010). Taken together, not only did the presence of 3DP PCL framework not have any detrimental impact on vessel formation, but it possibly may even enhance vascularisation within collagen-based scaffolds.

Indeed, several approaches to vascularise tracheal scaffolds have been investigated elsewhere in an attempt to accelerate neovascularisation of tracheal grafts, such as the application of angiogenic factors such as VEGF and human recombinant erythropoietin (hrEPO; 294,295). Although preliminary results showed promise in vascularising grafts, the maintenance of working concentrations of growth factors post-implantation has not yet been investigated (Tan et al., 2006). Furthermore, this approach relies on efficient host response and vascularisation may not occur in time to allow for graft survival. Another more widely-investigated approach has been pre-vascularising tracheal grafts through the means of implanting the graft into the patients forearm or omentum prior to implantation into the defect site (Lendlein et al., 2001; Delaere et al., 2019). The use of an initial heterotopic vascularisation period elsewhere in the body, prior to tracheal transfer, has been shown to generate a clinically significant blood supply (Kalathur et al., 2010; Kim et al., 2011; Udelsman et al., 2018; Li et al., 2019). However, such a strategy can take several weeks—or even months—to fully establish a vascular network and is therefore not adequate in emergency situations (Maughan et al., 2017). On the other hand, the means of pre-vascularising the reinforced composite CHyA-B scaffolds via the co-culture system, which we have shown can successfully vascularise *in vitro*, can establish a network of vasculature within the scaffolds to ensure success of graft acceptance into defect site within shorter time frames, as shown by previous success within our lab (McFadden et al., 2013). These mechanically reinforced pre-vascularised scaffolds can now be evaluated for their impact on enhancing epithelisation.

While this study has successfully achieved its objectives and developed a novel tubular scaffold for tracheal replacement, the lack of degradation data of our 3DP framework design, is a principal limitation. Although there are *in vivo* degradations studies on PCL that can provide an estimate 2–3 years degradation rate (Sun et al., 2006); the speed at which PCL breaks down is largely dependent on its surface area, geometry and molecular weight, all of which differ in studies. Completing a comprehensive degradation study would require a long-term animal study necessitating a large quantity of animals, which is unfortunately out the scope for the length of this research project. Reducing the number of animals required could be achieved by developing a non-invasive procedure to measure degradation within animals without requiring euthanasia, such as using micro-CT. Moreover, 3D-printed framework designs were mechanically assessed in a dry state due to the limitations of the testing machine, however, more appropriate assessment would be done under a hydrated state to mimic physiological conditions. Under hydration, PCL rate degradation would differ as the solution would enhance the hydrolytic breakdown of its ester bonds and thus allow for a more accurate measurement of each design's response to mechanical load. In particular, fatigue testing
performed in a hydrated state over a longer rate of cycles such as for
spinal implants (ISO 12189:2008) would provide vital information of the
scaffolds durability. Moreover, tracheal tissue is viscoelastic, as this study
has successfully established the foundation for a mechanically compliant
and flexible scaffold, future refinement of the scaffolds can look to
characterise its viscoelastic features.

5 Conclusion

Overall, we were successfully able to fabricate a composite tissue
engineered tracheal substitute that mimics the reported mechanical
properties of native tracheal tissue and confirmed the incorporation of
3DP PCL framework within the CHyA-B matrix did not impede on
vessel formation using an *in vitro* culture model. A battery of tests was
developed to characterise the 3DP designs by measuring mechanical
strength, robustness, and flexibility, all of which are integral tracheal
properties. This has enabled us to establish two main designs which
provide a range of mechanical properties close to animal models,
allowing us to streamline a design in the future when more data
becomes available on the exact mechanical properties of the trachea.
The use of PCL in tissue engineering, in particular tracheal
reconstruction, is becoming more widespread. Although PCL
exhibits biocompatibility it is less biocompatible than natural
polymers, and as previously discussed, tracheal scaffold attempts
with PCL alone have induced foreign body and inflammatory
reactions. However, the ease of printability and excellent mechanical
properties of PCL are substantial advantages, therefore we combined the
3DP PCL with freeze-dried CHyA-B matrices to create a composite
reinforced scaffold, which was confirmed to support the growth of cells
*in vitro*. The successful pre-vascularisation of our reinforced composite
CHyA-B scaffold lends itself resolving a major hurdle in the
advancement of tracheal scaffolds. By addressing both the
mechanical and physiological requirements of a tracheal scaffold,
this work has begun to pave the way for a new therapeutic option
to resolve the shortcomings of current treatment options, potentially
improving patient outcomes.

Data availability statement

The raw data supporting the conclusion of this article will be
made available by the authors, without undue reservation.

References

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Author contributions

TK—Experimental, Lead scientist in lab, manuscript draft
ML—Mechanical Testing and 3D printing contribution
LS—Mechanical Testing and 3D printing contribution
S-AC—Co-advisor on direction of work and overall study
performance FO'B—Co-advisor on direction of work and overall
study performance CO'L—Lead PI, idea generation, core study
concept, project direction, final manuscript editing and
submission. All authors listed have made a substantial, direct,
and intellectual contribution to the work and approved it for
publication.

Funding

This study was funded by the Science Foundation Ireland
Advanced Materials and Bioengineering Research (SFI-AMBER)
Centre (Grant 17/RC-PhD/3477).

Acknowledgments

We would like to acknowledge Dr Brenton Cavanagh and Dr
Ronaldo Jose Farias Correa Do Amaral, RCSI, for the respective
microscope training and CAM training provided to the first author.

Conflict of interest

The authors declare that the research was conducted in the
absence of any commercial or financial relationships that could be
construed as a potential conflict of interest.

Publisher's note

All claims expressed in this article are solely those of the authors
and do not necessarily represent those of their affiliated
organizations, or those of the publisher, the editors and the
reviewers. Any product that may be evaluated in this article, or
claim that may be made by its manufacturer, is not guaranteed or
endorsed by the publisher.

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OPEN ACCESS

EDITED BY

Vasil M. Garamus,
Helmholtz Centre for Materials and
Coastal Research (HZG), Germany

REVIEWED BY

Kai Wang,
Nankai University, China
Yang Zhu,
Zhejiang University, China

*CORRESPONDENCE

Juan Du,
✉ bairuochen12@163.com
Dian Chen,
✉ cdism221@163.com
Feiying Liu,
✉ liufeiying861218@163.com
Yang Dong,
✉ dongyang6405@163.com

¹These authors have contributed equally to this work

RECEIVED 16 March 2023

ACCEPTED 13 June 2023

PUBLISHED 22 June 2023

CITATION

Zhang H, Zhang Q, Du J, Zhu T, Chen D, Liu F and Dong Y (2023), Nanofibers with homogeneous heparin distribution and protracted release profile for vascular tissue engineering.

Front. Bioeng. Biotechnol. 11:1187914.
doi: 10.3389/fbioe.2023.1187914

COPYRIGHT

© 2023 Zhang, Zhang, Du, Zhu, Chen, Liu and Dong. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.

# Nanofibers with homogeneous heparin distribution and protracted release profile for vascular tissue engineering

Hongmei Zhang<sup>1,2†</sup>, Qilu Zhang<sup>3†</sup>, Juan Du<sup>2*</sup>, Tonghe Zhu<sup>1,2</sup>, Dian Chen<sup>4*</sup>, Feiying Liu<sup>5*</sup> and Yang Dong<sup>1*</sup>

<sup>1</sup>Department of Orthopedics Surgery, Shanghai Sixth People's Hospital Affiliated to Shanghai Jiao Tong University School of Medicine, Shanghai, China, <sup>2</sup>School of Chemistry and Chemical Engineering, Shanghai University of Engineering Science, Shanghai, China, <sup>3</sup>School of Textiles and Fashion, Shanghai University of Engineering Science, Shanghai, China, <sup>4</sup>Department of Cardiothoracic Surgery, Shanghai Children's Medical Center, Shanghai Jiao Tong University School of Medicine, Shanghai, China, <sup>5</sup>School of Biotechnology, East China University of Science and Technology, Shanghai, China

In clinic, controlling acute coagulation after small-diameter vessel grafts transplantation is considered a primary problem. The combination of heparin with high anticoagulant efficiency and polyurethane fiber with good compliance is a good choice for vascular materials. However, blending water-soluble heparin with fat-soluble poly (ester-ether-urethane) urea elastomer (PEEUU) uniformly and preparing nanofibers tubular grafts with uniform morphology is a huge challenge. In this research, we have compounded PEEUU with optimized constant concentration of heparin by homogeneous emulsion blending, then spun into the hybrid PEEUU/heparin nanofibers tubular graft (H-PHNF) for replacing rats' abdominal aorta *in situ* for comprehensive performance evaluation. The *in vitro* results demonstrated that H-PHNF was of uniform microstructure, moderate wettability, matched mechanical properties, reliable cytocompatibility, and strongest ability to promote endothelial growth. Replacement of resected abdominal artery with the H-PHNF in rat showed that the graft was capable of homogeneous hybrid heparin and significantly promoted the stabilization of vascular smooth muscle cells (VSMCs) as well as stabilizing the blood microenvironment. This research demonstrates the H-PHNF with substantial patency, indicating their potential for vascular tissue engineering.

KEYWORDS

heparin, electrospun vascular grafts, mechanical compliance, thrombosis, patency rate

## 1 Introduction

In recent years, the incidence of arteriovenous fistula and peripheral vascular diseases are increasing year by year, and vascular bypass transplantation is the most important treatment method (Martinez-Gonzalez et al., 2017; Caliskan et al., 2020). Therefore, there is an urgent need for tissue engineered vascular grafts, especially small-diameter vascular grafts. The main challenge to the limited application of small-diameter artificial blood vessels is that early thrombosis, lumen stenosis, intimal hyperplasia and low long-term patency rate (Tatterton et al., 2012; Seifu et al., 2013; Zhu et al., 2021a; Zhuang et al., 2021; Li et al., 2023). In addition, the mechanical compatibility of small-diameter vascular grafts should be the focus of attention (Zhu et al., 2020; Zhu et al., 2021a; Zhi et al., 2022). Therefore, to
achieve a long-term patency rate, the ideal small-diameter vascular grafts should have proper mechanical properties, outstanding blood compatibility and biocompatibility (Stahl et al., 2023).

Biomedical polymeric nanofibers present unique advantages and wide applications in the field of vascular tissue engineering due to three-dimensional biomimetic extracellular matrix (ECM) structure, adjustable physical and mechanical properties, stable biocompatibility as well as easy processing (Fathi Karkan et al., 2019; Nemati et al., 2019; Purushothaman et al., 2020). Poly (ester-ether-urethane) urea (PEEUU) is a biocompatible polyurethane derivative, which present excellent physical and mechanical properties and special microphase separation structure (Asadpour et al., 2018; Zhu et al., 2022). The soft and hard segments form the key surface structure of hydrophilic and hydrophobic alternating blocks in PEEUU, which not only conforms to Okano's hypothesis of the relationship between anticoagulant blood and material surface structure, but also has a similar structure as in human blood vessels (He et al., 2010). But pure PEEUU still cannot achieve the ideal blood compatibility and histocompatibility. Therefore, through continuous exploration, previous researchers have developed a series of methods or technologies to improve the anticoagulation of polyurethane, which included the improvement of the hydrophilicity of polyurethane, the constructing of ionic polyurethane surface, surface grafting or blend using bioactive macromolecules.

Among many anticoagulant units, poly (ethylene glycol) (PEG) is often selected as the coating component grafting on the surfaces of vascular lumen to reduce the adsorption rate of plasma protein and platelet adhesion (Tseng and Park, 1992; Altankov et al., 2000). Kim et al. reported a method for grafting end-sulfonated PEG on the surfaces of polyurethane membranes. The results of blood compatibility showed that the PEG-grafted membrane significantly reduced the adhesion of platelets and prolonged the APTT time (YoungKim et al., 2003). Zhu et al. reported a functionalized nanofibers vascular grafts with PEG and heparin via the sequential covalent grafting method. The grafted PEG and heparin synergistically decreased the water contact angle of nanofibers as well as increased the capability for preventing platelet deposition (Zhu et al., 2021b). Lee et al. reported a modification method of PEO-PPO-PEO triblock copolymer mixed with polyurethane and then formed a film by solvent volatilization. The results revealed that the antiplatelet adhesion of the modified polyurethane was improved and increased with the increase of the proportion of PEO chain segments in the block copolymer (Lee et al., 1998). The surface modification using zwitterion is also an alternative method to improve the anticoagulation of materials (Ukita et al., 2019; Wang et al., 2020; Zhu Tianyu et al., 2021). The modified surface can reduce the adhesion of platelets on the surface of the material through electrostatic repulsion between negative charges. Yuan et al. also reported a surface grafting technique based on plasma treatment. Firstly, polyurethane membrane was treated with $O_3$ plasma. Then, acrylic acid was grafted on the membrane. Finally, 1, 3-propyl sulfonolactone was grafted on the membrane by two methods. The blood compatibility test showed that the anticoagulation of polyurethane grafted with sulfonic acid anion group was significantly improved (Mokhtari and Kharazi, 2021). Among many bioactive macromolecules, heparin has excellent

anticoagulant properties and is the most commonly used anticoagulant. After more than 40 years of continuous development, the immobilization technology of heparin is showing more and more improvement in medical field. Although heparin presented obvious anti-coagulation and anti-hyperplasia effects, the existing immobilization methods, such as physical blending, ionic bonding, and covalent bonding by solution blending or solution grafting, cannot solve the problems of fast release, easy inactivation, and structural changes of heparin, which may be due to its very short half-life and easy dissolution in water rather than in organic solvents (Serrone et al., 2013; Kevin, 2015; Yang et al., 2022).

In this work, we propose a novel small diameter vascular graft to solve the above mentioned problems that the uneven dispersion of heparin in vascular material. Based on our previous research on nanofibers for sustained drug release and vascular tissue regeneration, the optimized concentration of heparin, was combined with synthetic PEEUU elastomer, following hybrid PEEUU/heparin nanofibers prepared by homogeneous emulsion blending and electrospinning for exploring blood vessel repair effect and mechanism. We evaluated the physicochemical properties, microscopic morphology, and cytocompatibility of hybrid PEEUU/heparin nanofibers, and then detected the activity of promoting tissue regeneration and inhibiting thrombosis and intimal hyperplasia. Finally, the abdominal aortic defect model of rats was used to evaluate the effect of hybrid PEEUU/heparin nanofibers tube.

# 2 Materials and methods

## 2.1 Materials

Poly (ester-ether-urethane)urea elastomer (PEEUU, Mw = 9.2 × 10⁴) were synthesized via a two-step solution polymerization in synthesis laboratory of functional polymer of Shanghai University of Engineering Science. Heparin sodium powder (150 U/mg) (Hep) were purchased from Sigma-Aldrich Trading Co., Ltd. (Shanghai, China). Human umbilical vein endothelial cells (HUVECs) for in vitro experiments were obtained from Shanghai Cell Bank of Chinese Academy of Sciences (Shanghai, China). Cell counting kit (CCK-8) was obtained from Sigma-Aldrich Trading Co., Ltd. (Shanghai, China). Dulbecco's modified eagle medium (DMEM), Fetal bovine serum (FBS), and antibiotic-antibacterial medicine (penicillin/streptomycin) were purchased from Hyclone Trading Co., Ltd. (Shanghai, China). Hexafluoroisopropanol (HFIP, purity ≥99.2%) was obtained from Shanghai Darui Fine Chemicals Co., Ltd. (Shanghai, China). Unless otherwise specified, all the above reagents were used directly. All the materials were used as received, except where mentioned otherwise.

## 2.2 Preparation of nanofibers tubes

To prepare hybrid PEEUU/heparin nanofibers tubular graft (H-PHNF), PEEUU was dissolved in 10 mL of HFIP to form a mixture with 8.0 w/v% concentration with stirring until clarified.
<table><caption>TABLE 1 Settings of PEEUU, heparin sodium dosage, and HFIP bulk in control and experimental groups, respectively.</caption><thead><tr><th>Samples</th><th>PEEUU (g)</th><th>Heparin sodium (mg)</th><th>HFIP (mL)</th></tr></thead><tbody><tr><td>PNF</td><td>0.8</td><td>0</td><td>10</td></tr><tr><td>B-PHNF</td><td>0.8</td><td>10</td><td>10</td></tr><tr><td>H-PHNF</td><td>0.8</td><td>10</td><td>10</td></tr></tbody></table>

0.01 g of heparin sodium (150 U/mg) was dissolved in three drops (about 0.2 mL) of distilled water until clarified, following addition of above-clarified PEEUU/HFIP mixture forming electrospinning solution with a relative mass fraction $W_{\text{Hep}}/W_{\text{PEEUU}}$ of 1.25% under vigorous stirring at room temperature for 72 h, respectively. Then, the 10 mL of the above prepared mixture solution were electrospun to generate nanofibers with a constant speed of 1.0 mL/h and a voltage of 12 kV.

Electrospinning solution for preparing PEEUU nanofibers (PNF) was prepared according to the following protocol: Synthetic 0.8 g PEEUU was dissolved in 10 mL HFIP formed a mixture with 8% concentration with stirring until clarified. To preparing the electrospinning solution for fabricating the blend PEEUU/heparin nanofibers (B-PHNF): Synthetic 0.8 g PEEUU was dissolved in 10 mL HFIP formed a mixture with 8% concentration. After stirring until clarified, quantitative heparin sodium powder (Hep) was added in the above-clarified mixture forming electrospinning solution with a relative mass fraction ($W_{\text{Hep}}/W_{\text{PEEUU}}$) of 1.25%.

The parameters to generate nanofibers showed in Table 1, respectively. The as-electrospun nanofibers were collected onto a stainless steel bar (2.0 mm diameter, 100.0 mm length, rotated at 200 rpm) or a flat aluminum foil board located 14 cm from the capillary to form nanofibers tubes or nanofibers mats, namely, PNF, B-PHNF, and H-PHNF, respectively. Then nanofibers tubes and nanofibers mats were vacuumed in a desiccator for 48 h to remove redissual HFIP.

## 2.3 Characterization and testing

The morphology and surface structure of nanofibers and nanofibers tubes were carried out using a scanning electron microscope (SEM, Phenom XL, Netherlands) operating with sputter gold plating for 35 s at 5 mA at an accelerating voltage of 10 kV. A contact angle measuring device (JC 2000D 2A, Shanghai Zhongchen Digital Technology Equipment Co., Ltd., China) was used to test the wettability of nanofibers tubes. To test this, 0.02 mL deionized water was added to the sample and three different positions of the sample were taken to measure the water contact angle and calculate the average value. Image-J (United States) was used to determine the diameter of nanofibers, inner diameter and wall thickness, and the pore diameter of nanofibers tubes.

The ethanol displacement method was applied to confirm the porosity of nanofibers tubes. The dried tubes were dipped in absolute ethanol, and the bubbles were removed. The porosity was determined by Eq. 1:

$$ \text{Porosity (\%)} = \frac{V_1 - V_3}{V_2 - V_3} \times 100\% \qquad (1) $$

where $V_1$ expresses the volume of known ethanol; $V_2$ stands for the bulk volume of the ethanol impregnated nanofibers tubes and ethanol; $V_3$ is the volume after removed the nanofibers tubes.

High-precision tensile testing machine (HY-025CS, Shanghai Hengyu Instrument Co., Ltd., China) with a transducer with a load range of 0–200 N was employed to testing the mechanical properties of nanofibers tubes in wet conditions at room temperature strictly according to ISO 7198: 1998. Each sample was cut into a cylindrical ring with length × inner diameter × wall thickness = 10.0 mm × 2.0 mm × 0.4 mm, then soaked in 0.01 M phosphate buffer (PBS, pH = 6.8) for 24 h. Finally, tensile tests were investigated at room temperature with a stretching speed of 1.0 mm/min. The specimens were extended until breaking under tensile force and the tensile stress-strain curves were recorded. Each test was repeated five times during mechanical analysis. The tensile strength, elongation at tensile strength, the representative modulus, and first order equation fitting of stress and strain within the magnification of calculating range were calculated according to the results of stress-strain.

## 2.4 Heparin density and sustained release tests in vitro

The presence of heparin in fibers were verified via toluidine blue staining. Briefly, a 0.005% toluidine blue solution was prepared in hydrochloric acid (0.01 M) containing 0.2% NaCl. The samples were incubated in the prepared toluidine blue solution for 12 h and dried at room temperature for further observation. The heparin density in fibers were verified via prepared toluidine blue staining assay. Electrospun tubes (2.0 mm inner diameter, 15.0 mm length) were immersed in 50 mL aqueous solution of toluidine blue [500 µmol/L toluidine blue solution dissolved in Milli Q water with a pH value of 10 (adjusted with NaOH)] and incubated in a vapor-bathing constant temperature vibrator at 37°C with a vibrating speed of 100 rpm for 12 h. Samples were then washed with dilute NaOH (pH = 10) five times followed by immersing in 50 mL aqueous solution of acetic acid with a concentration of 50% at 37°C with a constant vibrating speed of 100 rpm for 30 min. The OD value of toluidine blue released from fibers was read at 633 nm using a microplate reader (MK3, Thermo, United States).

The nanofiber samples for sustained release were placed in a constant temperature vibrator at 37°C (100 rpm). At the set time points, 3 mL of release solution was removed, while 3 mL of fresh PBS medium was added. Three test samples were set for each group. The cumulative release amount of heparin was calculated according to the standard heparin absorbance concentration at 260 nm. The percentage of accumulated release can be calculated by Eq. 2:

$$ \text{ARP (\%)} = \frac{C \times 30 + \sum W}{m \times R \times 1000} \times 100\% \qquad (2) $$

C-the concentration of heparin, µg/mL; $\sum W$-the mass of heparin accumulated release, µg; m-the mass of nanofibers tubes, mg; R-the percentage of drug within nanofiber.
## 2.5 Cell culture and blood compatibility tests in vitro

Human umbilical vein endothelial cells (HUVECs) were used to evaluate the activity of host cells, which were co-culture with PNF, B-PHNF, and H-PHNF nanofibers mats, respectively. HUVECs were obtained from Shanghai Cell Bank of Chinese Academy of Sciences (Shanghai, China) and cultured with growth medium consists of dulbecco's modified eagle medium (DMEM), 10% fetal bovine serum and 1% penicillin/streptomycin. The cell viability of HUVECs were tested by using the Cell Counting Kit-8 (CCK-8). The cells were cultured in PNF, B-PHNF, and H-PHNF nanofibers mats for 1 day, 3 days, and 5 days, respectively. Detailed procedures are available in the Supplementary Material.

Fresh blood, which was drawn from the marginal vein, and 3.2% sodium citrate solution in a volume ratio of 9:1 (v/v) were collected using a plastic vacuum blood collection tube (2.7 mL, Becton Dickinson, United States), containing 3.2% sodium citrate solution. All animal experimental protocols are in accordance with the policy of the Institutional Review Board for Human Investigations at Shanghai Sixth People's Hospital Affiliated to Shanghai Jiao Tong University School of Medicine. Detailed procedures are also available in the Supplementary Material (Mokhtari and Kharazi, 2021).

## 2.6 *In vivo* transplants to replace the abdominal aorta of rats

All rats were obtained from Shanghai Slaccas Experimental Animal Co., Ltd. (Shanghai, China) and all experimental schemes are in agreement with the requirements of the Institutional Animal Care and Use Committee (IACUC) of Shanghai Sixth People's Hospital Affiliated to Shanghai Jiao Tong University School of Medicine and accepted by IACUC. Ethical principles were followed throughout the experiment. All experimental plans were proceeded in conformity of the Animal Management Regulations of China (1988 and revised in 2001, Ministry of Science and Technology).

Sixty-three male Sprague-Dawley rats (age about 10 weeks; weighing approximately 300 g; 18 rats for standby application) were used as the abdominal aortic replacement models in the research. A 7.0 mm long defect of the abdominal aorta was created and replaced by nanofibers tubes (2.0 mm inner diameter, 7.0 mm length), and then the nanofibers tubes was end-to-end sutured to the abdominal aorta. Performance of the nanofibers tubes graft was evaluated on first, second, and fourth week after surgery in total of five animals per group for each time point.

At predetermined time points, rats were anesthetized by intraperitoneal injection of pentobarbital (3.5 mg/100 g; Nembutal). A color Doppler ultrasonography platform (GE LOGIQ 9, GE Medical Systems, United States) was used to acquire images to evaluate the patency rate of vessels. Then rats were sacrificed after injecting an overdose of isoprene barbiturate from the ear margin vein. Implanted nanofibers tubes were retrieved and analyzed by H&E staining, Masson's trichrome staining, Van Gieson, Saffron O, Von Kossa, and

immunofluorescent staining. Detailed procedures are available in the Supplementary Material.

## 2.7 Statistical analysis

Data were presented as mean ± standard deviation. All data were analyzed by one-way ANOVA with Tukey's *post hoc* tests. *p*-values <0.05 (*) were deemed to be statistically significant. * indicates *p* < 0.05, ** indicates *p* < 0.01, *** indicates *p* < 0.001.

## 3 Results and discussion

Synthetic vascular grafts perform satisfactorily as large-diameter (e.g., aortoiliac) arterial substitutes but commonly fail when employed in small-diameter applications (Wang et al., 2019). To date, clinically available expanded poly (tetrafluoro-ethylene) (ePTFE) or poly (ethylene terephthalate) (PET) vascular grafts are suboptimal for reconstructing small-diameter (inner diameter ≤6 mm) arteries, owing to thrombosis in early and restenosis in late stage (Lim et al., 2021; Roina et al., 2022). Moreover, ePTFE or PET vascular prostheses connected to human artery cannot swell when subject to the blood pressure owing to their poor elasticity, and thus also result in the decline of blood flow rate and the formation of thrombosis. Some previous research reported that polyurethanes were being applied in medical devices because they have the capability to tolerate contractile forces that originate during the cardiac cycle without undergoing plastic deformation or failure, and the capability to imitate the behaviors of different tissues (Rahimi and Mashak, 2013; Uscategui et al., 2018; Kelly Navas-GomezValero and Valero, 2020). More importantly, the mechanical compliance and other properties of polyurethane can be further improved by structural design and functional modification (Zhang Zhenyan et al., 2022). Therefore, polyurethane may be an optimal candidate.

Heparin is the most widely used and reliable anticoagulant that can enhance the patency and inhibit the thrombus. A lack of heparin leads to blood clotting, while an excess of heparin causes bleeding. (Kevin, 2015). Furthermore, the half-life of heparin is less than 1 h in circulation (Yang et al., 2022). Thus, maintaining localized concentration and homodisperse of heparin at the target site for long time is essential for implant revascularization. In this study, a hybrid electrospun BEPU/heparin nanofibers tubular graft compound with optimized constant concentration of heparin by homogeneous emulsion blending, were prepared for replacing rats' abdominal aorta *in situ* for comparing with nanofibers tubular graft, which were prepared via resolution after blending of solute.

### 3.1 Microstructure and mechanical properties

We have compounded PEEUU/HFIP solution with heparin/ distilled water solution, and then processed it into clarified PEEUU/ HFIP/heparin/distilled water mixture according to the solvent compatibility of HFIP and water. PEEUU/heparin could be feasibly dissolved in HFIP/distilled water mixture to form a
<figure><img src="image_129.png" /></figure>

homogeneous hybrid solution for random electrospinning and collected by a 316 L rolling stick collector to obtain tubular grafts (Figure 1A). Similarly, control samples were also prepared in this way, except for the different electrospinning solution.

The morphology of prepared tubular grafts are shown in Figures 1B,C. The SEM images of the lumen surface of PNF, B-PHNF, and H-PHNF, in Figure 1B, showing its nonwoven structure formed by random stacking of nanofibers with the average diameter of 607 ± 31 nm, 769 ± 55 nm, 655 ± 49 nm, respectively, (Figure 1D). The diameter of both PNF and H-PHNF group were finespun but evenly distributed, while agglomeration appeared in B-PHNF group. Correspondingly, the pore sizes of both PNF and H-PHNF group are smaller than those of B-PHNF groups, and the porosity is larger than that of B-PHNF group (Figures 1D–F).

Moreover, as shown in Figure 1C, the electrospun grafts could maintain its tubular shape after being peeled from the 316 L rolling stick collector. The statistical data of inner diameter showed that the average diameter of PNF, B-PHNF, and H-PHNF groups were 1.854 ± 0.152, 1.719 ± 0.189, and 1.756 ± 0.172 mm as well as the average wall thickness of PNF, B-PHNF, and H-PHNF were 283 ± 12, 241 ± 23, and 267 ± 17 µm, respectively (Figures 1G,H). Compared with B-PHNF group, the inner diameter sizes and wall thickness of both PNF and H-PHNF groups are larger than those of B-PHNF groups. We attributed these appearance to the fibers size and structure of grafts, and the distribution of heparin in fibers.

In the design of small-diameter artificial blood vessel, the first consideration should be the mechanical properties (Zhu et al., 2020). Due to its small diameter and thin wall, small-diameter artificial blood vessels are not strong enough to resist external forces.
<figure><img src="image_130.png" /></figure>

Therefore, the designed small-diameter artificial blood vessels should be able to resist certain pressure in the body without deformation. Moreover, the artificial blood vessels must have a burst pressure similar to or higher than that of human blood vessels, which have a burst pressure of more than 1700 mmHg.

We have measured the mechanical properties of nanofibers tubular grafts in wet conditions at room temperature strictly according to ISO 7198: 1998 using the uniaxial radial high-precision tensile testing machine as the schematic illustration in Figures 2A,B. As shown in Figure 2B, the zone of breaks of tubular graft is located in the middle of the sample, which means this fracture test is effective. Representative radial stress-strain curves and maximum tensile strength of PNF, B-PHNF, and H-PHNF are presented in Figures 2C,D, respectively. The results showed that the maximum tensile strength of PNF, B-PHNF, and H-PHNF are 5.096 ± 0.310 MPa, 4.596 ± 0.201 MPa, and 4.916 ± 0.251 MPa,

respectively, which are significantly larger than 1.365 ± 0.211 MPa of rats' abdominal aorta.

Natural blood vessels are viscoelastic bodies with viscoelastic characteristics such as creep and stress relaxation, which can be treated as elastic materials and have anisotropic nonlinear stress-strain relationship after sufficient preconditioning. As in Figures 2E–G, the representative radial stress-strain curves, representative 0–20% modulus, first order equation fitting of stress and strain within the magnification of calculating range of 0–20% modulus were calculated and fitted according to the dates in Figure 3E, respectively. It is clear that the modulus within the magnification of calculating range of 0–20% strain of H-PHNF significant increased compare to rats’ abdominal aorta (Figure 2F). The increase of modulus also proves that H-PHNF have good anti-deformation performance after implantation in vivo. In short, three electrospun tubular grafts showed reliable resiliency, while H-PHNF
<figure><img src="image_131.png" /><figcaption>FIGURE 3</figcaption></figure>

(A) The digital photos of PNF, B-PHNF, and H-PHNF, respectively; (B) Toluidine blue staining digital photos of PNF, B-PHNF, and H-PHNF in wet states, respectively; (C) Toluidine blue extract used acetic acid as treating agent; (D) Quantitative data of adsorption assays indicate in the amounts of toluidine blue in nanofibers tubular grafts; (E) Water contact angle at the 1, 3, and 5 s time point; (F) *In vitro* releases of heparin from B-PHNF and H-PHNF in pH = 6.8 PBS, respectively; (G, H) Higuchi equation fitting of cumulative release rate and release time of B-PHNF and H-PHNF, respectively. (Data are representatives of independent experiments and all data are given as mean ± SD, *n* = 5; \*p < 0.05, \*\*p < 0.01).

exhibited better initial elasticity, which was more suitable for application as a prosthetic blood vessel.

As shown in Figures 2H–J, the obtained burst pressure substantially reached those of both artery and saphenous vein which is the gold standard for bypass operation, whereas the suture retention strength was a little smaller than PNF but still quite larger than the rats’ abdominal aorta (around 0.293 MPa) for surgical operation.

Compliance is defined as the volume change in response to changes in blood pressure, and constitutes an important parameter that influences the short- and long-term patency of artificial blood vessels (Stahl et al., 2023). Especially in the early stage of transplantation, the mismatch of mechanical compliance between vascular grafts and autologous blood vessels is more likely to cause acute thrombosis at the anastomosis site. The compliance of PNF, B-PHNF, H-PHNF, and

rats' abdominal aorta were evaluated, respectively, as shown in Figure 2K. The compliance of both PNF and H-PHNF were larger than that of the other two groups. The compliance of H-PHNF is better than B-PHNF as well as rats' abdominal aorta. PEEUU/heparin tubular graft through a combination of homogeneous emulsion blending technology and electrospinning technology exhibit prominent stable mechanical compliance due to the homogeneous dispersion of heparin in the electrospun fibers.

## 3.2 Surface wettability and heparin sustained release properties

We obtained tubular grafts with a length of 2.0 cm and an inner diameter of 2.0 mm (Figure 3A). The presence of as hybrid heparin
was examined by toluidine blue staining and analyses exhibited color change from white to dark blue. Images of toluidine blue staining demonstrated that H-PHNF received more dye after staining (Figure 3B). This was in line with the quantitative study that the heparin-end groups had much higher absorbance than the other groups (Figures 3C,D). The results of water contact angle measurements are exhibited in Figure 3E. PNF showed its hydrophobic nature as the water drop well stood on its surface after it was dropped down and its contact angle was measured about 113° at the 1 s time point. When heparin molecule blended in the electrospun PEEUU fibers, the obtained B-PHNF exhibits higher hydrophilicity (the water contact angle is 74° at the 1 s time point), while the water contact angle of H-PHNF is 102° at the 1 s time point. We believe that the above phenomenon is not only closely related to the microstructure of the fiber, but also to the dispersion of the heparin molecules. Specifically, due to poor dispersion, more heparin molecules gathered on the surface of PEEUU fibers in B-PHNF, and these highly hydrophilic heparin molecules increased the roughness of the surface of PEEUU fibers. In general, the hydrophilicity of B-PHNF was stronger than that of H-PHNF.

In the process of drug release, the properties of drug release depend on complex physical and chemical phenomena (Zhu et al., 2021b). The factors controlling the drug release rate are not only the three-dimensional structure of the grafts, the interaction between molecules, the degradation rate of the material, the solubility of the drug, but also closely related to the dispersion of the drug in the polymer matrix. Figure 3F shows the cumulative drug release curve of heparin in B-PHNF and H-PHNF, where the X-axis is the release time, and the ordinate is the percentage of cumulative release in the total drug load. The release curves of heparin in B-PHNF and H-PHNF with similar drug concentration were compared. The release of heparin from both B-PHNF and H-PHNF mainly includes three stages: the initial drug release stage (0–24 h), controlled release stage (24 h–5 days), and late accelerated release stage (5 days~). At the initial release phase, the released heparin mainly comes from the surface of the PEEUU fiber. Most of this heparin, which exposed to the surface of the fiber, is very vulnerable to the surrounding tissue fluid wash and rapid release. It is obvious that heparin in both B-PHNF and H-PHNF exhibited obvious initial release, and the cumulative release rate at 3 h reaches 12.69% and 9.84%, respectively. On the 14th day, the final cumulative release of heparin in both B-PHNF and H-PHNF reached 91.26% and 71.67%, respectively, which shows a clear difference. In short, heparin in H-PHNF was released at the slower rate than B-PHNF after 24 h, which presumably achieve long-term inhibiting thrombus and intimal hyperplasia effect. Moreover, according to the time square root equation proposed by Higuchi, the least square regression analysis was carried out for the release time $t^{1/2}$ and the cumulative release amount $M_t/M(%)$. The fitting equations were $M_t/M(%) = 4.9907t^{1/2} + 3.9893$ and $M_t/M(%) = 3.5958t^{1/2} + 5.5162$, and the fitting correlation coefficients were 0.9562 and 0.9574, respectively. Figures 3G,H show that the release of heparin at the initial stage of B-PHNF and H-PHNF conforms to the Higuchi equation, and diffusion is the main mechanism of heparin release.

## 3.3 Hemocompatibility in vitro

When the material comes into contact with blood, it causes platelet adhesion and activation of the clotting system. Since platelet and coagulation system work together to produce coagulation, the interaction between the two should be carefully considered in the design of blood compatible materials (Zhu et al., 2021b). After the contact between the isolated venous blood and the material, coagulation factors and the endogenous coagulation system are activated, and finally fibrin is generated and blood coagulation occurs. The time experienced in this stage is called coagulation time. Coagulation time can generally reflect the total coagulation ability of blood after contact with materials. After the whole blood incubation, the absorbance of the supernatant was measured at 540 nm, indicating the number of remaining erythrocytes, as shown in Figure 4A. Compared with the other two groups, the H-PHNF had the maximum absorbance at any time point. Under the same experimental conditions, the clotting time of both PNF and B-PHNF were significantly shorter, indicating that H-PHNF has anticoagulant properties, which may be due to its outstanding hydrophobic performance according to the contact angle results (Figure 3A).

All the clotting factors involved in endogenous clotting are supplied by plasma. When plasma comes into contact with the material, the surface charge activates the clotting factors in the plasma. In the presence of $Ca^{2+}$, these clotting factors bind to $Ca^{2+}$ to form prothrombin complexes, which activate and convert to active thrombin in the presence of thrombin and $Ca^{2+}$. Subsequently, soluble fibrinogen in plasma is converted to insoluble fibrin by thrombin and $Ca^{2+}$, which eventually leads to clotting. Plasma recalcium curve is a method used to characterize the endogenous coagulation system. Recalcium time refers to the time required for plasma coagulation after the removal of calcium source and the addition of $Ca^{2+}$. The detailed anticoagulant mechanism is that Heparin forms a complex by binding with antithrombin III, which accelerates the inactivation of coagulation factors and thus inhibits the formation of prothrombin kinase. In addition, heparin calcium also acts on already formed prothrombin kinases, thus providing a stronger anticoagulant effect. Figure 4B shows the plasma recalcium kinetic curves of PNF, B-PHNF, H-PHNF, positive control group, and negative control group. The positive control has the fastest clot formation time at 15 min, while the clot formation time of PNF was 20 min. Notably, B-PHNF, H-PHNF consistently kept a low absorbance and closed to the negative control without emerging the inflection point. These results indicate that the grafts complexed with heparin can significantly inhibit endogenous coagulation activation and improve the anticoagulation performance of the grafts.

When the material comes into contact with red blood cells, the degree of damage to red blood cells can be characterized by the hemolysis rate. The smaller the hemolysis rate of the material, the better the blood compatibility of materials. The hemolysis rate test results of PNF, B-PHNF, and H-PHNF samples are shown in Figure 4C. Figure 4C showed that the hemolysis rates of the three grafts were all less than 5% as specified in the ISO10993-4 standard, indicating that the three grafts were non-hemolysis materials with good blood compatibility, non-toxicity, and no obvious damage to erythrocytes. Moreover, both the unheparined
<figure><img src="image_132.png" /></figure>

<figure><img src="image_133.png" /></figure>

<figure><img src="image_134.png" /></figure>

<figure><img src="image_135.png" /><figcaption>FIGURE 4</figcaption></figure>

(A) Whole blood clotting time; (B) Plasma recalcification time; (C) Quantification of relative hemolysis rate; (D) Platelet deposition determined by lactate dehydrogenase assay. (For plasma recalcification time test, TCPs exposed to PPP with and without CaCl₂ were used as positive control and negative control, respectively; For hemolysis test, water and 0.9% normal saline (NS) serve as positive and negative groups, respectively; Data are representatives of independent experiments and all data are given as mean ± SD, n = 5; \*p < 0.05, \*\*p < 0.01).

and heparin-modified grafts had hemolysis rates of approximately 3%, indicating that heparinization did not affect the hemolytic properties of grafts.

The adhesion and activation of platelets can activate clotting factors and promote the formation of fibrin, which eventually leads to the formation of thrombus (Jordan et al., 2016). Therefore, the resistance against adhesion of platelets on the surface of materials is one of the important markers of the antithrombotic properties of materials. To compare the influence of compound heparin on the number of adherent platelets, the results of quantitative detection of the number of adherent platelets using the lactate dehydrogenase (LDH) kit are shown in Figure 4D. The number of platelet adhesion per unit area on the surface of the material is shown in Figure 4D. The platelets with obvious adhesion are on the surface of PNF, while the number of platelet adhesion is significantly reduced on both B-PHNF and H-PHNF surfaces due to the anticoagulation of heparin. In addition, there was no significant difference between the B-PHNF and H-PHNF groups, indicating that no matter which technique was used to combine heparin and PEEUU, the spatial conformation of heparin would not change significantly, let alone affect the anticoagulant effect of heparin in the short term.

## 3.4 Cytocompatibility in vitro

The inner layer of natural arterioles is rich in endothelial cells, which can resist thrombosis and regulate the transmission of signals and substances inside and outside the blood vessels. As an

anticoagulant, heparin can avoid the clotting reaction induced by biomaterials in vivo as well as inhibits bacteria and cellulose adhesion. Introducing the biological function of heparin into polymer materials to achieve complete endothelization of the intima of tissue engineered vascular stent as soon as possible is an important content of the research on tissue engineered artificial blood vessels.

The proliferation of HUVECs grown onto PNF, B-PHNF, and H-PHNF after 3-day culture were evaluated, respectively (Figures 5A–C). It can be found from the fluorescence microscopy images in Figure 5A, the number of HUVECs in both B-PHNF group and H-PHNF groups were more than PNF group. As shown in Figures 5A,B, although the absorbance of the B-PHNF group was significantly higher than that of the H-PHNF group, the morphology of HUVECs in the H-PHNF group was significantly better than that in the B-PHNF group, and a large number of pseudopods were appeared. Figure 5C shows the growth morphology of HUVECs after culturing on the grafts for 3 days. It can be clearly observed that the number of endothelial cells adhered to B-PHNF is significantly higher than that of other groups, while HUVECs adhered to H-PHNF possessed abundant lamellipodia and filopodia, which indicates that the grafts loaded with heparin have almost no cytotoxicity and can support the adhesion and proliferation of endothelial cells. Furthermore, we selected two marker genes, eNOS and VEGF, to determine the functionality of HUVECs using reverse transcription polymerase chain reaction (RT-PCR) after 3 days of culture. H-PHNF markedly promoted HUVECs-related gene expression in vitro (Figures 5D,E).
<figure><img src="image_136.png" /><figcaption>FIGURE 5</figcaption></figure>

(A) Fluorescence microscopy images of HUVECs grown onto PNF, B-PHNF, and H-PHNF with labeling of cytoskeleton (red) and nucleus (blue) after 3-day culture, respectively; (B) CCK-8 assay of the proliferation viability of HUVECs cultured on PNF, B-PHNF, and H-PHNF for 1, 3, and 5 days culture, respectively; (C) SEM images of HUVECs grown onto PNF, B-PHNF, and H-PHNF after 3-day culture, respectively; (D, E) HUVECs-related gene expression of eNOS and VEGF after 7 days of culture, respectively. (Data are representatives of independent experiments and all data are given as mean ± SD, *n* = 5; \*p < 0.05).

<figure><img src="image_137.png" /><figcaption>FIGURE 6</figcaption></figure>

(A) Schematic diagram of nanofibers tubular grafts implanted into rats' abdominal aorta via end-to-end anastomosis; (B) Microscopical images of the surfaces and cross-sections of the retrieved B-PHNF, H-PHNF, and its neighboring native blood vessel after implantation for 4 weeks; (C) Patency rate statistics for each of electrospun tubular grafts at each time point (n = 6).
<figure><img src="image_138.png" /></figure>

The H-PHNF showed a synergistic promotive effect of heparin and matching microstructure on rapid endothelization and upregulated the expression of vascular endothelial cell-related genes.

## 3.5 Patency after abdominal aorta implantation of rats

We prepared grafts via homogeneous emulsion blending and electrospinning, followed by replacing isometric left abdominal aorta of SD rats with B-PHNF and H-PHNF by end-to-end anastomosis, respectively (Figure 6A). Nanofibers tubular grafts were retrieved for overall assessment 4 weeks after surgery. Macroscopically, the retrieved tubular graft maintained its white color and opaque texture, which can be easily distinguished from the native blood vessel that had semi-transparent appearance. More importantly, the surface of the endovascular lumen of unobstructed tubular grafts were smooth and there were no thrombus (Figure 6B). Tubular graft and native blood vessel had comparable diameters and fused well to form integrated tissue at the suture site (Figure 6B). The patency rate were detected after 1 week, 2 weeks, and 4 weeks

transplantation, respectively, and the results are shown in Figure 6C. After 1 week transplantation, all the samples in both B-PHNF and H-PHNF groups remained unobstructed, while only 60% patency rate of PNF group was noted. With the increase of transplantation time, three tube of B-PHNF were blocked in the second week. In the fourth week, one tube was blocked in the H-PHNF group, and five tubes of both PNF and B-PHNF were blocked, respectively. Compared with the other groups, the H-PHNF had a longer-lasting release of heparin, which had leed to the obvious effect of inhibiting thrombus formation. This is mainly due to the technology that promotes more uniform mixing of PEEUU and heparin. Ultimately, H-PHNF exhibited stronger ability to blocks the rapid release of heparin molecules than B-PHNF.

## 3.6 Histological assessments after implantation *in situ*

As shown in Figure 7, the cross-sections of B-PHNF and H-PHNF in proximal end, midportion, and telecentric end after
<figure><img src="image_139.png" /></figure>

4 weeks of implantation were stained against CD31 and α-SMA antibodies and immunofluorescent imaged to demonstrate the presence of endothelial cells and smooth muscle cells, respectively. Both B-PHNF and H-PHNF had the apparent expressions of α-SMA after implantation for 4 weeks. Compared with B-PHNF, the regenerative SMCs layer in the lumen of H-PHNF was more analogous in SMCs layer thickness, no matter in proximal end, midportion, and telecentric end, which is mainly due to the uniform distribution of heparin in H-PHNF.

The intimal layer had positive expression of endothelial cells and smooth muscle cells. The internal wall of both B-PHNF and H-PHNF lumen were completely covered by an endothelial layer composed of monolayered and uniformly distributed endothelial cells (Figure 7). No thrombosis occurred within the lumen of grafts after 4 weeks. Taken together, these data indicated that hybridized heparin contributed to enhanced anticoagulant capacity, conferred rapid endothelization, and limited intimal hyperplasia.

The results of H&E and Masson's trichrome staining at proximal end, midpoint, and telecentric end of transverse section of B-PHNF and H-PHNF for 4 weeks implantation are shown in Figure 8. H&E and Masson's trichrome staining were used to observe the lumen area, neo-tissue formation and wall thickness of the patent grafts (Zhang Yan et al., 2022). The values of lumen diameter of proximal end, midpoint, and telecentric end in H-PHNF group were 1.98 ± 0.07 mm, 1.98 ± 0.10 mm, and 1.98 ± 0.09 mm. There was no big difference among these groups, while they were smaller than that of

the B-PHNF group (1.92 ± 0.11 mm, 1.96 ± 0.08 mm, and 1.93 ± 0.13 mm for proximal end, midpoint, and telecentric end, respectively) (Figure 8A). The quantitative data is shown in Table 2. In addition, neo-tissue was observed on the luminal surface no matter in B-PHNF group or H-PHNF group. Moreover, there was discontinuous neo-tissue coverage detectable at the midpoint in B-PHNF group, while continuous neo-tissue coverage detectable at in H-PHNF group. Both H&E and Masson's trichrome staining showed that the structure of the implanted tubular grafts were composed of a large amount of radially circumaxial oriented collagen in both inner layer and outer layer, which may be secreted by the smooth muscle layer. Therefore, we speculate that the main source of SMCs encapsulating tubular grafts is these smooth muscle tissues are mainly from the autovascular tissue migrated from the anastomosis (Figure 8B). All in all, the migration of SMCs also indicates that tubular grafts may have vascular functions such as relaxation and contraction in the later stage.

Van Gieson staining showed that large amounts of elastic fibers were secreted in both the inner wall and outer wall of the artificial blood vessel, and elastic fibers was arranged along the circumference of the tubular grafts (Figure 9A). Elastin is an important guarantee of vascular elasticity and compliance, while elastic fibers mainly provides mechanical support. When the tubular grafts were implanted in the body to guide the orientation regeneration of smooth muscle cells, elastic fibers-oriented deposition was also
<table><caption>TABLE 2 The lumen diameter at proximal end, midpoint, and telecentric end of B-PHNF and H-PHNF for 4 weeks implantation, respectively.</caption><thead><tr><th>Samples</th><th>Proximal end (mm)</th><th>Midpoint (mm)</th><th>Telecentric end (mm)</th></tr></thead><tbody><tr><td>B-PHNF</td><td>1.92 ± 0.11</td><td>1.96 ± 0.08</td><td>1.93 ± 0.13</td></tr><tr><td>H-PHNF</td><td>1.98 ± 0.07</td><td>1.98 ± 0.10</td><td>1.98 ± 0.09</td></tr></tbody></table>

<figure><img src="image_140.png" /></figure>

achieved, which enhanced the radial mechanics of the tubular grafts and prevented the occurrence of aneurysms.

Another important substance in the natural extracellular matrix is mucopolysaccharide (GAG), a negatively charged straight-chain carbohydrate composed alternately of hexose carboxylic acids and hexose amine groups. GAG can not only form bond and regulate growth factors and cytokines, but also inhibit the decomposition of protease, and plays a positive role in the adhesion, migration,

proliferation and differentiation of endothelial cells. Safranin O staining showed proteoglycan distribution on both outer and inner surfaces of tubular grafts (Figure 9B). Similarly, GAG was evenly distributed in the lumen of H-PHNF with a approximate thickness of GAG deposition, no matter proximal end, midpoint, or telecentric end, which is mainly due to the uniform distribution of heparin in H-PHNF, and that heparin is one of the main factors that stimulate the secretion of GAG in host cells.
Calcific degeneration remains a major obstacle facing the translation of tissue-engineered vascular grafts for arterial repair. The reasons that cause calcification are complex, including the elasticity of the material, the acid degradation products, and the hardness of the material. To check whether calcification had occurred after 4 weeks implantation, Von Kossa staining of transverse section of the explanted tubular graft was performed (Figure 9C). The results showed that no calcification was detected in the explanted tubular graft. Some previous studies have shown that heparin can inhibit the differentiation of vascular smooth muscle into osteoblasts in the environment as well as inhibit intracellular mineral deposition. In addition, the nanofibers with heparin homogeneous distributions promote sustained ECM secretion by host cells, which inhibits calcification due to the degradation of collagen.

## 4 Conclusion

In summary, anti-acute thrombosis cues of anticoagulant molecule heparin were incorporated into a nanofibers tubular graft (H-PHNF) of great compliance match with the native blood vessel by homogeneous emulsion blending and electrospinning. Structure and component modification confer to moderate wettability, matched mechanical and sustained drug release profile, and reliable blood compatibility of the obtained H-PHNF. H-PHNF showed prominent ability to promote endothelial cells growth and maintained patency without acute thrombosis formation in abdominal aorta implantation of rats in situ. This strategy may bring cell-free small-diameter synthetic vascular grafts toward clinical application.

## Data availability statement

The original contributions presented in the study are included in the article/Supplementary Material, further inquiries can be directed to the corresponding authors.

## Ethics statement

The animal study was reviewed and approved by the all rats were obtained from Shanghai Slaccas Experimental Animal Co., Ltd. (Shanghai, China) and all experimental schemes are in agreement with the requirements of the Institutional Animal Care and Use

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Committee (IACUC) of Shanghai Sixth People's Hospital Affiliated to Shanghai Jiao Tong University School of Medicine and accepted by IACUC.

## Author contributions

HZ: experiment, data analysis, manuscript writing; QZ: methodology, manuscript writing; JD: supervision, data analysis, manuscript writing and editing; TZ: Supervision; DC: supervision, methodology; FL: supervision, manuscript review and editing; YD: supervision, manuscript review and editing.

## Acknowledgments

The authors sincerely appreciate projects sponsored by the National Natural Science Foundation of China (81902186), Shanghai Rising-Star Program (20QC1401300), Technology Support Project of Science and Technology Commission of Shanghai Municipality (22S31904700).

## Conflict of interest

The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

## Publisher's note

All claims expressed in this article are solely those of the authors and do not necessarily represent those of their affiliated organizations, or those of the publisher, the editors and the reviewers. Any product that may be evaluated in this article, or claim that may be made by its manufacturer, is not guaranteed or endorsed by the publisher.

## Supplementary material

The Supplementary Material for this article can be found online at: https://www.frontiersin.org/articles/10.3389/fbioe.2023.1187914/full#supplementary-material

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