Tracheal tissue regeneration

12

Tracheal tissue regeneration

F. A c o c e l l a and S. B r i z z o l a, Università degli Studi di Milano, Italy

Abstract: Tracheal circumferential defects involving more than half of the tracheal wall still represent an unsolved problem. Several studies have developed different methods to help repair cartilage and improve healing but a suitable tracheal reconstruction or replacement has not been achieved yet. Novel bioengineering technologies seem to be the new answer to this serious problem. This chapter briefly describes the fundamentals of the anatomical and physiological tracheal functions and provides a review of trachea tissue engineering. Then it describes the project and development by means of electrospinning of a biodegradable tubular tracheal scaffold with an in vitro and in vivo preliminary experimental approach. Key words: animal model, Degrapol®, electrospinning, neovascularisation, tissue engineered trachea (TET).

12.1

Anatomy of the trachea and main pathologies of surgical concern

12.1.1 Tracheal anatomy The trachea is a single cartilaginous and membranous conduit crucial for the ventilation and the clearance of upper air-way secretions. Part of the trachea lies in the cervical region while a small part of it runs under the sternal notch through the thoracic cavity. The structure of the trachea in the adult human male averages 11.8 cm in length (range 10–13 cm) and 1.6–2.4 cm in width from the lower border of the cricoid cartilage to the top of the carinal spur. The trachea is nearly cylindrical, being flattened posteriorly and composed of 18–22 C-shaped rings of hyaline cartilage (Fig. 12.1). The posterior part of the trachea consists of a muscular structure (pars membranacea) that stretches under inflow and outflow pressure. At the spur level the trachea divides into two main bronchi (right and left) that provide continuity to the respiratory system. The right main bronchus continues more vertically, whereas the left is always more horizontal with respect to the trachea, while in infants the two bronchi lie more transversely. As a base rule in infants more than half of the trachea is found within the neck, whilst in an adult half of the organ is positioned above the sternal notch. This proportion varies during flexion and extensions of the neck caused by the tracheal movement. During childhood the cartilage is thinner and eventually more compressible laterally than in 242 © Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 242

6/6/11 6:50:16 PM

Tracheal tissue regeneration Oesophagus

Tracheal ring

243

Larynx

Trachea

Bronchi

Left

Right

12.1 Drawing of the human trachea.

adults. The cross-sectional configuration of the trachea may be markedly altered with increasing age, particularly in the presence of chronic obstructive lung disease which could cause different shape deformities with potential clinical relevance. The blood supply to the organ is crucial for tracheal surgical therapy and it is essentially based on three tracheoesophageal branches coming from the inferior thyroid artery. The first branch supplies the lower cervical trachea. The second and the third supply the middle and the upper sections respectively. The bronchial arteries originating from the aorta supply the carina and the lowest part of the trachea. Starting from the vessels reaching the trachea a lateral longitudinal anastomosis is seen. This is crucial for the entire blood supply of the organ. From here several anterior transverse intercartilaginous arteries run deeply in the trachea anastomosing with the contralateral arteries at the midline. This vascular network reaches the inner part of the trachea and submucosa, and forms the sub-mucosal plexus that provides sole nourishment to the cartilage and is an important source of blood for the membranous wall and mucosa. Ultrastructurally the trachea is formed by hyaline cartilage (forming the tracheal ring), smooth muscle cells (forming the pars membranacea) and, going deeper, the sub-mucosal layer which houses the vascular plexus, as discussed above. The luminal surface is lined by the respiratory mucosa made by a pseudostratified columnar epithelium with cilia and goblet cells. These two different kinds of cells are essential to produce mucus for moisturising the airway and excluding foreign particles using vibratile cilia (Fig. 12.2).

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 243

6/6/11 6:50:16 PM

244

Electrospinning for tissue regeneration 147.9 µm

A 192 µm

211.9 µm

B

C

12.2 Ultrastructural section of the tracheal wall showing hyaline cartilage (A), sub-mucosal vascular plexus (B) and mucosal layer (C)

12.1.2 Pathologies of surgical concern About half of the trachea could be successfully resected and reconstructed by end-to-end anastomosis. Grillo (2002) This is the basic concept of tracheal surgery. On the other hand pathologies involving more than half of the trachea need tracheal substitutes, biological or synthetic, but several clinical situations lead to organ amputation and creation of a tracheostomy with consequential aphonic voicelessness and potential airway infection. In general it is possible to declare that all the pathologies involving the trachea could result in tracheal resection and substitution owing to their extension of disease. As described by Grillo (2003) tracheal diseases can be split into two main categories: malignant and benign. Tracheal tumours are mainly malignant. Primary tumors are squamous cell carcinomas and ‘cylindromas’ (adenoid cystic carcinomas), while secondary tracheal neoplasia are bronchogenic carcinomas of the carina and thyroid carcinomas (owing to their anatomical relationship). Among benign lesions we can distinguish congenital disease (tracheal agenesia, stenosis, esophago-tracheal fistula and laringo-tracheal cleft) and acquired disease (post-intubation lesions and strictures, subglottic laryngeal stenosis, tracheoesophageal strictures and tracheal trauma).

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 244

6/6/11 6:50:16 PM

Tracheal tissue regeneration

12.2

245

Tissue engineered trachea (TET)

12.2.1 Tissue engineering and tracheal replacement The management of tracheal pathology, such as stenosis or cancer, often requires a tracheal reconstruction. Primary end-to-end anastomosis after resection is the method of choice and can usually be performed successfully for defects of up to 50% in length (Kon, 1983). Unfortunately, there are cases in which primary anastomosis is not possible, such as after extensive burns, trauma, tumour resection, or post-intubation injuries. Additionally, the treatment of congenital tracheal atresia or stenosis can be hindered by the lack of sufficient tissue for surgical reconstruction, as the length of trachea involved may be extensive. Over the last 60 years, multiple approaches to tracheal reconstruction have been attempted both clinically and experimentally, including the use of prosthetics, autografts, homografts and allografts. Prostheses such as Dacron polyurethane mesh, polytetrafluoroethylene, polypropylene mesh, silicone rubber, and even glass tubes (Jacobs, 1988) have been tried and have often been fraught with complications such as infection, extrusion and stenosis (Table 12.1). Autogenous and alloplastic tissues have also been used from sources such as fascia, skin, bone, periosteum, cartilage, perichondrium, tracheal allografts, muscle, oesophagus, pericardium, dura mater and the small bowel (Fonkalsrud, 1971; Cohen, 1985; Har-El, 1989; Letang, 1990). In any case the main problem for all these grafts remains the development of late stenosis of the prosthesis. Because of these difficulties, investigators are still in search of the ideal tracheal replacement material, which is able to satisfy the requirements of a tracheal prosthesis model and is safe and durable in practical use. First, as Belsey (1950) iterated in the 1950s, the tracheal substitute should be a laterally rigid and longitudinally flexible tube even if the latter requirement has proved to be desirable, but not essential (Grillo, 1965). The tracheal conduit must further be initially airtight and become integrated into adjacent tissues, so that chronic inflammation, granulation tissue, infection and erosion do not occur. Also we cannot forget that the need for an immunosuppressive regime is unfavourable for different reasons but especially in the frequently extensive tracheal cancer that requires a transplant. Furthermore, for a method to be practically considered, the technique of construction or insertion of the conduit must be surgically straightforward and the results must be predictably successful. Finally the materials for tracheal replacement must be biocompatible, non-toxic, nonimmunogenic, non-carcinogenic, must not dislocate or erode over time, should ideally provide or facilitate epithelial resurfacing, should avoid stenosis or late buckling, resist bacterial colonisation, avoid accumulation

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 245

6/6/11 6:50:16 PM

246

Electrospinning for tissue regeneration

Table 12.1 General approaches used for tracheal replacement

Type

Complications

Synthetic materials Solid

Steel Glass Polyethylene Silicone Teflon Polyurethane

No incorporation and epithelisation

Porous

Steel wire Titanium Collagen PTFE Polyurethane Dacron Teflon

Scar tissue formation Lack of epithelisation

Bioprosthesis Etherologous tissues Cadaver trachea Aortic homograft Autogenous tissues

Fascia Aorta Tracheal wall Cartilage Derma Pericardium Periosteum Buccal mucosa Dura mater Bone strips

Vascularised Intercostal muscle autogenous tissues Diaphragm Oesophagus

Scar tissue formation Deformation, contraction, calcification Malacia and degeneraton Need for foreign material support

Need for foreign material support Requirement for major surgery

of secretions, and must be permanent constructions (Jackson, 1950; Scherer, 1986). Since 1988, tissue engineering of cartilage has been successfully pursued by a number of groups. Osada (1994) and his group carried out a tracheal replacement in ten mongrel dogs using a long knitted Dacron® tube with inner silicone rubber coating seeded with autologous fibroblasts after static culture. The animals long term survival was satisfactory and the results encouraging. During the same year, Vacanti (1994) and associates were able to create new hyaline cartilage in athymic mice by using isolated bovine chondrocytes and biodegradable suture material (polyglactin 910 and polyglycolic acid, (PGA)), which served as a temporary scaffold to which cells attached until they created their own matrix.

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 246

6/6/11 6:50:16 PM

Tracheal tissue regeneration

247

One of the main limitations in this approach is the specific mechanical requirement placed on the trachea. The trachea must maintain flexibility in the longitudinal direction to allow for free movement of the head, whilst maintaining the rigidity necessary to prevent collapse of the trachea during breathing. This is accomplished in native tissue by presence of cartilaginous rings, and is not adequately modelled by a cartilaginous tube. For these reasons Kojima (2002), designed another study to evaluate the ability of autologous tissue-engineered cartilage with a helical shape to provide the structural component of a functional tracheal replacement (Fig. 12.3(a)).

(a)

(b)

12.3 (a) Helical template fabricated with a silicone mold-making kit. (b) The chondrocyte-seeded matrix was placed in the grooves of the template (arrow) and the entire template was wrapped with the fibroblast-seeded mesh.

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 247

6/6/11 6:50:16 PM

248

Electrospinning for tissue regeneration

Chondrocytes and fibroblasts were seeded onto separate nonwoven meshes of PGA fibres, 14 mm in diameter (Davis and Geck, Danbury, CT, USA). The chondrocyte-seeded mesh was placed in the grooves of a helical template made from Silastic ERTV mold-making kit (Dow Corning) and then covered with the fibroblast-seeded mesh (Fig. 12.3(b)). These implants were placed either in a subcutaneous pocket in the nude rat or in the neck of a sheep. Sheep tissue-engineered tracheas were harvested from the neck at 8 weeks and anastomosed into a 5 cm defect in the sheep trachea. Gross morphology and tissue morphology were similar to that of native tracheas and histology revealed the presence of mature cartilage surrounded by connective tissue. Other biomaterials have also been used in combination with chondrocytes to produce high quality tissue-engineered cartilage. These biomaterials include calcium alginate gels, collagen gels, fibrin glue, and agarose gels, among others. Another study by Kojima (2003a) evaluated the feasibility of creating engineered tracheal equivalents grown in the shape of cylindrical cartilaginous structures using sheep tracheal and nasal septum cartilage-derived chondrocytes. Tracheal and nasal chondrocytes were separately seeded onto PGA matrices and these cell–polymer constructs were then implanted subcutaneously in nude mice for 8 weeks. This approach demonstrated that the cell yield and the property of the resulting engineered cartilages were very similar for both tracheal and nasal chondrocytes (Fig. 12.4).

(a)

(b)

(c)

(d)

12.4 Appearances of tracheal chondrocytes (a) and (c) and nasal chondrocytes (b) and (d) seeded on non-woven mesh of polyglycolic acid (PGA) fibres (Davis and Geck, Danbury, CT) for 8 weeks to create tissue engineered trachea (TET).

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 248

6/6/11 6:50:17 PM

Tracheal tissue regeneration

249



In a follow-up study (Kojima, 2003b), the same group evaluated the feasibility of producing a composite engineered tracheal equivalent lining its lumen with nasal epithelial cells. At first chondrocyte suspensions were seeded onto a matrix of PGA. Cell–polymer constructs were wrapped around silicon tubes and cultured in vitro for one week, followed by implanting into subcutaneous pockets on the backs of nude mice. After 6 weeks the epithelial cells were suspended in a hydrogel and injected into the embedded cartilaginous cylinders following removal of the silicon tube which until then had been used as an internal support for the engineered construct. Implants were harvested after 4 weeks and the morphology of implants resembled that of native sheep trachea while the histology showed the presence of mature cartilage and formation of a pseudo-stratified columnar epithelium. Other polymers were used in subsequent studies with the same success, including a nonwoven mesh of PGA and PGA/poly-l-lactic acid copolymers. Recently, Ruszymah (2005) and colleagues undertook a study to reconstruct the trachea with human nasal septum chondrocytes by using a combination of biodegradable hydrogel and non-biodegradable high-density polyethylene (HDP) as the internal predetermined shape scaffold. After 8 weeks of in vivo implantation, the TET constructs were harvested. The macroscopic appearance of the TET constructs demonstrated that the HDP constructs were 80–90% covered with yellowish glistening cartilage-like tissue without any sign of inflammation. Okumus (2005) and colleagues designed an axial biosynthetic prefabricated flap to reconstruct the circumferential tracheal defects in ten rabbits. The inner mucosal lining was substituted by hairless epithelium obtained from the proximal ear, the tracheal cartilage by polypropylene mesh and the tracheal adventitia by lateral thoracic fascia to ensure a good vascular supply. Then the epithelial graft, polypropylene mesh and lateral thoracic fascia were put into a tube around a silicone catheter and placed into the cervical subcutaneous area for 2 weeks. A silicone catheter was taken out and prefabricated neotrachea adapted to the defect formed in native trachea and anastomised. The constructs were evaluated after 4 weeks by radiological, macroscopical and histological examination. In high level active research there are essentially three emerging concepts: the use and development of dynamic in vitro culture systems (bioreactors), the use and development of microfabrication technologies to create vascularised tissues and organs, and the search for and use of an appropriate multipotent, undifferentiated stem cell in tissue engineering. Producing a dynamic in vitro microenvironment for tissues may be an important aspect in guiding the formation of tissue with certain structural and functional characteristics. The in vitro modulation of chondrogenesis by dynamic cell seeding and bioreactors has been investigated. Vunjak-Novakovic (1999) and colleagues found that hydrodynamic conditions in convective-flow tissue culture bioreactors can

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 249

6/6/11 6:50:17 PM

250

Electrospinning for tissue regeneration

modulate the composition (matrix components), morphology, mechanical properties and electromechanical function of tissue engineered cartilage. These initial results are encouraging for the development of a TET replacement in humans and tissue engineering has emerged as a rapidly expanding approach for addressing the organ shortage problem.

12.2.2 First transplant of a human tissue engineered tracheal prosthesis Moving from their own previous and successful research Macchiarini (2008) have made a tubular tracheal human transplant. A 30-year-old female patient with end-stage airway disease underwent the complete resection of the left main bronchus with a bioengineered human trachea replacement. This was achieved by using autologous epithelial cells and chondrocytes that were isolated from biopsies and subsequently cultivated in vitro. The confluent cells were then seeded onto a matrix consisting of decellularised tracheal segment retrieved from a 51-year-old female transplant donor. The chondrocytes were seeded onto the external surface whilst the epithelial cells were seeded onto the internal side with the same density of 1 ¥ 106 cells per ml. The newly seeded tracheal segment was incubated in a bioreactor for 96 hours to control cell proliferation and matrix deposition. After this time, the patient underwent reconstructive surgery with positive results. The post-operative course was uneventful and at 14 days, 1 month, 2 months and 3 months the graft appeared to be healthy. This is a highly important step for the production of a tissue engineered tracheal substitute and gives great hope for the future of the airway reconstruction.

12.2.3 Tracheal cartilage, structure and biomechanics Owing to the structural and functional complexity of the organ, the various tissue components must be considered independently to realise a feasible tracheal analogous substitute. It is clear that the cartilaginous component of the tracheal rings will play a central role in the development of a functional tissue, being resistant to the inspiratory collapse and, at the same time, characterised by adequate flexibility. Tracheal cartilage, a type of hyaline cartilage, is a composite material in which the collagen fibres, essentially type II, are immersed in a hydrated proteoglycan matrix (aggrecan, versican, hyaluronane). Its main function is to keep the airway wall open despite intrathoracic pressure differences during breathing that would otherwise cause it to collapse and limit air flow. Moreover the cartilaginous rings are crucial in determining the dimension and compliance of the trachea; indeed, the rings limit prolapse of the pars membranacea, when the transmural pressure is positive, or its invagination,

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 250

6/6/11 6:50:17 PM

Tracheal tissue regeneration

251

when the transmural pressure is negative. Human tracheal cartilage consists of a collagen-rich ablumenal superficial zone, a proteoglycan-rich core and a collagen-rich lumenal superficial zone. At each surface there is a gradual transition between the superficial zone of the cartilage and a collagenous perichondrium. Chondrocytes are in lacunae surrounded by extracellular matrix and in the superficial, collagen-rich zones are closer together and flattened parallel to the plane of the nearest cartilage surface. Collagen fibrils in the superficial zones are oriented in the plane of the cartilage surface whilst in deeper layers of the cartilage, collagen fibrils are oriented less regularly and the proteoglycan content decreases (Fig. 12.5). The biomechanical properties of the tracheal cartilage are a consequence of its structural features. Any biomechanical change in airway cartilage could influence the mechanics of maximal expiratory flow and cough. Age-related changes in biomechanical properties and biochemical composition of airway cartilage could influence the airway dynamics owing to the shortening of airway smooth muscle. (a)

(b)

(c)

(d)

12.5 Scanning electron microscopy images (SEM) of the different tracheal cartilage zones: (a) Luminal surface is to the left with chondrocyte lacunae visible in an abundant regular layering of the matrix (Scale bar: 40 µm). (b) Higher power view of the boxed area from (a) showing the luminal superficial zone. Rows of collagen fibrils (arrows) are oriented perpendicular to one another (Scale bar: 10 µm). (c) SEM of the peripheral abluminal side with perichondrium on the left (Scale bar: 100 µm). (d) Higher power view of the boxed area from (c) showing perpendicular fibres deposited on this side (Scale bar: 10 µm).

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 251

6/6/11 6:50:17 PM

252

Electrospinning for tissue regeneration

The equilibrium tensile modulus of airway cartilage may be most directly relevant to normal breathing, where cartilage provides a relatively constant load to the trachealis muscle and maintains airway wall stiffness over a range of airway diameters that are produced by changes in the tone of the trachealis muscle occurring over minutes to hours. Furthermore, the equilibrium tensile modulus is likely to be independent of proteoglycans in this tissue, as has been shown for articular cartilage (Schimdt, 1990). In contrast, in a forced expiration or cough, transmural pressure changes occur in less than a second and the viscoelastic properties of the airway cartilage may be most important in maintaining airway calibre. Under either a sustained force or a transient force, the resistance of tracheal cartilage to bending depends on the tensile modulus in layers furthest from the neutral axis (i.e. the ablumenal superficial zone) and the compressive modulus of the central zone of the cartilage. The tensile modulus is primarily a function of collagen fibril organisation and the compressive modulus is a function of aggregating proteoglycans that are abundant in the central zone. There is substantial experimental data in the literature (Rains, 1992) regarding the Young’s modulus value (E) of the tracheal cartilage which differs widely depending on age-related individual changes or cartilage portion and layer creating structural inhomogeneity, whilst the Poisson modulus value (n) is generally constant (Table 12.2).

12.3

Electrospun biodegradable tubular tracheal scaffold

12.3.1 Scaffolds and tissue engineering This section describes the development of a biodegradable microstructured tubular scaffold created by electrospinning to mimic the rabbit native trachea shape and functional characteristics and its experimental evaluation both in vitro and in vivo. The architecture of an engineered tissue substitute plays an important role in modulating tissue growth. The primary function of a scaffold is tissue conduction and, therefore, it must allow cell attachment, migration onto or within the scaffold, cell proliferation and cell differentiation. It must also provide an environment where the cells can maintain their phenotype Table 12.2 Mechanical features of the human tracheal cartilage (tensile properties). E is Young’s modulus value; n is Poisson modulus value Age(years)

E (MPa)

n

References

80–81 36–74 17–81 17–81

16.67 2.5–7.7 1.8–15 4.6–13.6

0.3 0.3 0.3 0.3

Begis (1988) Lambert (1991) Rains (1992) Roberts (1991)

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 252

6/6/11 6:50:17 PM

Tracheal tissue regeneration

253

and synthesise required proteins and molecules. Scaffold characteristics include high porosity, high surface area, structural strength and specific three-dimensional (3D) shape. Although various structures of engineered tissue scaffolds have been developed for tracheal replacement, the goal of producing a clinically useful tissue scaffold is still far from being realised. An ideal tissue engineered scaffold should be mechanically stable and able to function biologically in the implant site (Thomson, 1997). Mechanical stability is dependent primarily on the selection of the biomaterial, the architectural design of the scaffold and the cell–material interactions. The ultimate goal of the scaffold design is the production of an ideal structure that can replace the natural extracellular matrix (ECM) until host cells can repopulate and synthesise a new natural matrix. To achieve this goal the scaffold material must be carefully selected and the scaffold architecture must be designed to ensure that the seeded cells are biocompatible with the engineered scaffold. The surface chemistry of a tissue engineered scaffold is dependent upon the type of the material, ranging from natural biopolymers to synthetic polymers. The most commonly used natural biopolymers include demineralised bone matrix (Dahlberg, 1991), agarose and collagen (Watt, 1988), hyaluronan (Allemann, 2001), basement membrane (Ponticiello, 2000) or alginate (Bonaventure, 1994). Synthetic polymers that are used include degradable polyesters, such as PGA and polylactic acid (PLA) and their copolymers, poly(d,l-lactide-co-glycolide) (PLGA) (Agrawal, 2001). These biodegradable polymers have a long history of clinical use and currently are used in various tissue engineering applications. Furthermore the tendency of host tissue to form a fibrous capsule around an implant is an important challenge for scaffold biocompatibility. The encapsulation of porous implants depends partially on their material architectural features. There are in general two architectures of porous biomaterial implants: those made of interconnected fibres (fibro-porous meshes) and those made of interconnected pores (porous meshes). Softtissue response is sensitive to the geometric features of them both (Davilla, 1968; Jansen, 1992). Basically materials with small pores sizes (<50–80 mm) will experience less tissue ingrowth and greater encapsulation than those with larger pore sizes (> 80–100 mm). Sanders (2000, 2005) and his group demonstrated the fundamental role of the fibre diameter and pore dimension on cell adhesion and proliferation on an electrospun fibro-porous mesh (Fig. 12.6). It is clear, however, that tight control of fibre diameter is important. Some studies reported in the literature (Paqu ay, 1996) suggest that the reason why large-diameter fibre meshes become encapsulated is related to capsule formation around individual fibres within the mesh. Capsules from adjacent fibres may eventually interact to form a single capsule around the entire perimeter of the implant. It is only with large fibre spacing that interaction between

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 253

6/6/11 6:50:17 PM

254

Electrospinning for tissue regeneration 70

Implant with a well-defined capsule at the perimeter (%) Encapsulated fibres in implant cross-sections (%)

Encapsulation (%)

60 50 40 30 20 10 0

0

10

20

30 40 50 Fibre spacing (µm)

60

70

12.6 Encapsulation dependence on fibres spacing. Only the 6 µm spacing meshes showed encapsulation around the perimeter. All meshes showed some encapsulation of individual fibres (Sanders, 2005).

70 µm

25 µm

Fibre Chondrocyte

12.7 Bidimensional model showing ideal space in intra-fibres (30/80 µm) and ideal fibre diameters (6/25 µm) for fibre-porous meshes to be more suitable for adhesion, proliferation and differentation of cells.

capsules around adjacent fibres and the consequent perimetric encapsulation would be avoided. With small-diameter fibres, where most of the fibres are not encapsulated, interaction between adjacent capsules is reduced and the minimum spacing between fibres necessary to avoid encapsulation at the perimeter is reduced. Thus, according to what is reported in the literature the ideal fibre-porous mesh with 6/25 mm fibres diameter and an inter-fibre space of 30/80 mm could promote cells adhesion, viability, proliferation and also exchange of nutrients (Fig. 12.7). The electrospinning technique is effective in generating very small diameter fibres ranging from 0.05 mm to 5 mm which are useful in different

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 254

6/6/11 6:50:17 PM

Tracheal tissue regeneration

255

applications such as separation membranes, wound dressing materials or artificial blood vessels (Doshi, 1995; Hohman, 2001). In our studies, we chose the DegraPol® polymer because it combines the advantages of traditional polyesters with high processability and marked elasticity properties; its use has been investigated for a long time, showing promising results especially in the fields of cartilage tissue engineering, with appropriate cellular attachment, growth and proliferation. As described by Saad (1996–1999) Degrapol® has been successfully tested under different forms (i.e. foam) both in vitro and in vivo using different kinds of cells.

12.3.2 DegraPol®: a degradable block polyesterurethane for tissue engineering DegraPol® is a degradable block polyesterurethane, consisting of crystallisable blocks of poly((R)-3-hydroxybutyric acid)-diol and blocks of poly(ecaprolactone-co-glycolide)-diol linked by a diisocyanate. Being a block copolymer, DegraPol® combines the advantages of traditional polyesters with high ductility and marked elasticity properties. The use of DegraPol® foams and porous membranes as scaffolds for tissue engineering has been investigated for a long time, showing promising results especially in the fields of cartilage (Raimondi, 2004), trachea (Yang, 2003) and smooth muscle (Danielsson, 2006) tissue engineering. There is significant evidence of DegraPol® demonstrating in vitro and in vivo biocompatibility properties. Saad described in his studies the response of in vitro cultured fibroblasts, macrophages and osteoblasts to the crystalline domain of the material. He reported that phagocytosis of the crystalline segments cause dose-dependent cell activation, cell damage and cell death in macrophages but not in fibroblasts. These studies moreover indicated that DegraPol® exhibits good cell compatibility and does not induce cytotoxic effects in osteoblasts and chondrocytes. In another study, Saad (1997) also performed subcutaneous implantations in rats of polyesterurethanes samples processed into films (5 mm in diameter and 150 mm in thickness) by compression molding; the results of these in vivo experiments showed that two months after implantation the thickness of the capsule was lower than 30 mm. Borkenhagen (1998) described in vivo performances of DegraPol® as a tubular structure used as a nerve guidance channel, showing that the inflammatory reaction associated with polymer degradation does not interfere with the nerve regeneration process. Recently (Riboldi, 2005), DegraPol® scaffolds were manufactured by electrospinning in the novel form of microfibrous membranes with fibres about 10 mm in diameter and a fibre-to-fibre distance of about 10 mm. This advance, together with the above cited findings about the polymer biocompatibility, suggests that DegraPol® holds promise for use as a scaffold for tracheal cartilage tissue engineering.

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 255

6/6/11 6:50:18 PM

256

Electrospinning for tissue regeneration

12.3.3 DegraPol® characteristics The molecular weight of the electrospun polymer used was 60 765 Da measured by gel permeation chromatography (GPC) and it was chemically composed of hard segments (40%) and soft segments (60%), providing the material with specific long-term degradation behaviour in water at 37.5°C, as shown in Fig. 12.8.

12.3.4 DegraPol® electrospun tubular scaffold A small tubular structure was obtained by collecting fibres on a solid, cylindrical, rotating iron collector (~300 rpm) with a 5 mm diameter, fixed to a drill and positioned 20 cm from the pipette (Fig. 12.9). A chloroform Degrapol® solution (Merk KGaA) was initially heated to a maximum temperature of 40°C preventing the polymeric solid particles from degrading during melting. Remarkable changes in the collected fibres were observed when varying the electrospinning and/or solution parameters: ∑

∑

Solution flow-rate: larger fibre diameters were observed with increasing flow rate as shown in Fig. 12.10. The mean fibre diameter obtained with a flow rate of 1.19 ml h–1, 2.38 ml h–1 and 3.57 ml h–1 were 10 mm, 20 mm and 25 mm, respectively. Temperature: Fig. 12.11 shows scanning electron microscope (SEM) 100000 90000 Molecular weight (Da)

80000 70000 60000 Mp

50000 40000

Mw

30000 20000 10000 0 0

20

40

60

80 100 Time (days)

120

140

160

180

12.8 DegraPol® degradation in buffered aqueous solution at 37°C. After 4 weeks of hydration, the average Mp of the polymer has dropped to approximately 45% of the initial value. At 12 weeks, this value has decreased further, approaching 25% of the starting value. Within 25 weeks, the average Mp value is below 10 000 Da (Mp is the peak average molecular weight, Mw is the weight average molecular weight).

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 256

6/6/11 6:50:18 PM

Tracheal tissue regeneration

257

1 cm

12.9 Collector spindle, consisting of a 5 mm diameter metal cylinder, fixed to a drill. A thin layer (white) of fibres can be seen on the spindle. 50 µm

(a)

100 µm

(b)

50 µm

(c) ®

12.10 Three chloroform Degrapol solutions (Merk KGaA) were electrospun at different flow rates: the fibre diameter increased with increasing flow rate. (a) 1.19 ml h–1, (b) 2.38 ml h–1, (c) 3.57 ml h–1 (Scale bar = 50 µm). 50 µm

(a)

50 µm

50 µm

(b)

(c)

12.11 Three scaffold surfaces obtained under different thermic conditions, resulted in increasing interconnected fibres at higher temperatures. (a) 25°C, (b) 20°C, (c) pre-dipping in liquid nitrogen (Scale bars = 50 µm (a, c) and 100 µm (b)).

images of reticulate fibres produced under different environmental thermic conditions. The electrospun scaffold porosity, resulting from interconnections and melting of fibres, appears to vary according to the

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 257

6/6/11 6:50:18 PM

258

∑

Electrospinning for tissue regeneration

environmental temperature (T), that is porosity decreases with higher temperatures. The first photograph (SEM) was taken in the summer, at an ambient temperature of approx. 25°C and 60% humidity; the middle one was taken in the winter, at an ambient temperature of approx. 20°C and 40% humidity. In the third image, the collector spindle had been pre-dipped in liquid nitrogen (–170°C), so that the fibres are fully disconnected and porosity is remarkably increased. Concentration: the solution concentration plays a major role in the fulfilment of reticulate fibres characterised by specific size, intrafibre spaces and porosity. The solution viscosity increases with higher polymer content, resulting in an increased mean fibre diameter and porosity and decreased fibre interconnections as described below. A solution concentration of 25–27% meets our requirements perfectly.

Various tests led to reticulate fibres with a diameter ranging from 10 to 25 mm, intrafibre spaces of 30–70 mm and good porosity, resulting from poor fibre interconnections. Figure 12.12 indicates that the reticulate fibre obtained under the conditions listed in Table 12.3 is similar to the model described in Fig. 12.7.

12.4

Scaffold fulfilment

12.4.1 Selection of the scaffold profile In mammals, the tracheal architecture and profile create a fixed cross-sectional area of the trachea during extreme bending caused by body, head and neck

22 µm

80 µm

30 µm

12.12 SEM Image of an electrospun scaffold section with fibre spacings measured according to the standards specified in literature.

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 258

6/6/11 6:50:18 PM

Tracheal tissue regeneration

259

Table 12.3 Operational conditions and solution properties utilised in the production of electrospun reticulate fibres Variables

Values

Solution flow rate Temperature Solution concentration

3.57 ml h–1 –170°C 27% (w/w)

The collector spindle was pre-dipped in liquid nitrogen prior to covering with fibres.

movements, preventing the obstruction of its inner lumen and the resulting respiratory arrest. This particular physiological behaviour derives from the regular distribution of cartilage rings along the trachea and the bronchi and prompted us to create a scaffold with this specific structural characteristic. Therefore, we studied three different profile models: spiral, ringed and toothed, with identical functional properties mimicking as near as possible the anatomical peculiarities and functional necessities of the mammalian trachea, as shown in Fig. 12.13. The goal was to identify which of these three profile models resulted in the best shape and function and could then be tested in a preclinical animal model.

12.4.2 Development of the spiral profile In the case of bending stress, the spiral profile prevents kinking of the structure, as the coils collide filling the intra-ring spaces, whilst leaving the inner lumen unchanged. First, a very thin tubular structure with a diameter of 0.8 mm was spiral-coiled around the collector wire (0.5 mm diameter) and then covered with a further layer of fibres (6/25 mm diameter) providing stability and consistency for the structure. The spiral scaffold (Fig. 12.14) obtained, shows all the design characteristics depicted in Fig. 12.13. The real issue for this type of electrospun structure was represented by nonhomogeneous fibre deposition, mainly in the proximity of the coils, resulting in a local high-porosity size (> 80–100 mm) structure that is unfavourable to the survival of cells. The same problem occurred to the internal spiral-rolled pipe, although it would have favoured the passage of growth factors during cell culturing. Finally, assembly of the structure caused many problems, such as manually extracting the small pipe from the wire or coiling the pipe around the spindle at a specific distance.

12.4.3 Development of the ringed profile A 1 mm thick electrospun tubular scaffold with a 5 mm diameter was initially utilised for the development of the ringed scaffold. After this scaffold had been covered with cyclohexane and slowly dipped in dry ice for refrigeration,

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 259

6/6/11 6:50:18 PM

260

Electrospinning for tissue regeneration 29

6.6

8

5

1.8

9

Spiral

5

7

2

6

1.25

1

Ringed

5

9

2

1

Toothed

12.13 Three scaffold models and relevant measurements (mm) investigated for tracheal scaffolds. The diameter of the inner lumen is 5 mm for all models.

it was fixed perpendicularly to the blade for the next process. A cryostat (Shandom ‘Microtome’ 5030, Great Britain) operating under low temperature ambient conditions (–20°C), has been used to dissect the tubular scaffold into rings 2 mm in length. The same 5 mm collector spindle was slipped into the rings spaced at 1 mm intervals and coated with an extra layer of electrospun fibres (6/25 mm diameter). The problems reported in this case coincide with those of the spiral scaffold. The inconsistency of porosity near the rings and assembly problems were major issues. The ringed scaffold is shown in Fig. 12.15 and Fig. 12.16

12.4.4 Development of the toothed profile The starting conditions of the electrospinning equipment and the solution (Table 12.3) are unchanged, including the type of collector spindle with a

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 260

6/6/11 6:50:18 PM

Tracheal tissue regeneration

261

5 mm

(a)

5 mm

(b)

12.14 Trachea-scaffold type 1 made by electrospinning: spiral profile. External structure (a) and manual bending (b) demonstrating the structure kink resistance.

12.15 Ringed profile. Image showing the longitudinal section of the scaffold level with the ring.

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 261

6/6/11 6:50:19 PM

262

Electrospinning for tissue regeneration 5 mm

(a)

5 mm

(b)

12.16 Trachea-scaffold type 2 made by electrospinning: ringed profile. External structure (a) and manual bending (b) demonstrating the structure kink resistance.

5 mm diameter. First, the collector spindle was dipped in liquid nitrogen for a few seconds and immediately processed for electrospun fibres deposition. After approximately 30 min a layer of fibres 2 mm thick (6/25 mm diameter) has been obtained and we proceeded with the mechanical lathe processing of the scaffold. In this case, a 1-mm thick blade for metal cutting has been used running for 2 mm along the longitudinal section of the scaffold (Fig. 12.13) and leaving a 0.5 mm layer of fibres on the spindle. All the problems that related the inconsistency of the porosity in both previous scaffolds, were overcome by means of a layer of fibres obtained in a single process. Moreover, this method was carried out in controlled automated operational steps. Figure 12.17 shows the toothed scaffold.

12.4.5 Inner layer In order to obtain an inner homogeneous surface that would be free of tissue overproductions, inclinations and narrowing of the surface, we decided to

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 262

6/6/11 6:50:19 PM

Tracheal tissue regeneration

263

5 mm

(a)

5 mm

(b)

12.17 Trachea-scaffold type 3 made by electrospinning: toothed profile. External structure (a) and manual bending (b) demonstrating the structure kink resistance.

install a very thin electrospun microporous layer (approximately 100 mm thick with fibres of 20/25 mm diameter) in the inner lumen, enabling permeability of gas, growth factors and cells nutrients, and, at the same time, avoiding cell proliferation and cartilage tissue growth, which would otherwise result in the occlusion of the inner lumen. This microporous layer was obtained by electrospinning a different solution composed of Degrapol® (16% w/w) dissolved in chloroform (90%), with the addition of methanol (10%). This particular layer has been obtained by processing the collector spindle for electrospun fibres deposition for approx. 1 min (ambient temperature of approx. 25˚C) prior to refrigeration and coating it with the final layer of fibres (Section 12.4.4). Figure 12.18 shows the SEM image of the obtained inner layer.

12.5

In vitro and in vivo evaluation of the cell and tissue response

As previously described, one of the major approaches in tissue engineering includes the seeding and culture of cells on a 3D scaffold prior to its

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 263

6/6/11 6:50:19 PM

264

Electrospinning for tissue regeneration

100 mm

12.18 SEM image of the inner degrapol® electrospun microporous layer made from a solution of polymer dissolved in chloroform and methanol (90:10 mix).

implantation. Other fundamental requirements concerning the scaffold geometry, such as porosity, pore interconnectivity or pore size to ensure development of an in vitro functional tissue after cell seeding attachment and migration, are of great importance. Moreover it has been demonstrated that during expansion in the monolayer, chondrocytes de-differentiate, assuming a more flattened appearance and producing type I instead of type II collagen; however, the chondrocytic phenotype may be rescued by transferring cells to a 3D culture system (Benya, 1982; Bonaventure, 1994; Jakob, 2001). Our initial in vitro experiments investigated the static culture of chondrocytes on the electrospun microstructure scaffold, in order to determine whether our scaffold provided a good 3D environment for cell adhesion and proliferation and whether the chondrocytic phenotype was preserved. Later we investigated the effect and reliability of dynamic conditioning on the development of in vitro tracheal engineered tissue for a period of 14 days. Lastly, we determined which tissue harvesting technique was the easiest, safest and most minimally invasive and what was the optimal source of cartilage cells. For this reason we investigated the isolation and expansion of rabbit chondrocytes from different anatomical sites to determine an ideal cell concentration for cell seeding.

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 264

6/6/11 6:50:19 PM

Tracheal tissue regeneration

265

12.5.1 In vitro static culture Briefly autologous chondrocytes were harvested from xifoid, chondrocostal and femural cartilage of a New Zealand White rabbit, digested in 0.3% collagenase II, and the cells serially passed. After the expansion phase (6–8 weeks) the chondrocytes were seeded onto the external surface of a half-cut scaffold with a seeding density of 25 ¥ 106 cell/ml, and the cells–polymer construct maintained during in vitro static culture (Fig. 12.19). Five days after static cell seeding, SEM images showed an even distribution of chondrocytes throughout the scaffold, suggesting appropriate exchange of gas and nutrients (Fig. 12.20(a) and (b)).

12.5.2 In vitro dynamic culture We designed a culture model system with an ad hoc bioreactor to allow better colonisation of the electrospun tracheal scaffold that favours the transport of nutrients from the medium to the adhering cells through strict control of culture parameters. Briefly, scaffolds were installed on the bioreactor and chondrocyte suspensions were seeded onto the external scaffold surface with gentle rotation in order to cover the whole circumference of the scaffold. The seeded scaffolds were transferred to the incubator (37°C, 5% CO2) and maintained in culture medium for about 2 hours to enhance cell attachment. After this time, the bioreactor was turned on gradually to reach a rotation speed of 5 rpm, moving up to 7 rpm after a further two hours until the end of the culture period. The scaffold was imaged by environmental scanning electron microscopy

1 mm

Seeding area

12.19 SEM micrograph showing 5 days of static culture for chondrocytes on the seeding area surface of the Degrapol® toothed profile matrix (magnification 12x).

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 265

6/6/11 6:50:20 PM

266

Electrospinning for tissue regeneration

20 µm

(a)

20 µm

(b)

12.20 SEM micrographs demonstrating adhesion and infiltration of chondrocytes on the scaffold surface (a) and within intrafibre spaces in the 3D scaffold (b) after 5 days of static culture.

(ESEM) following culture for 14 days in the bioreactor. The images demonstrated the adherence of the cell on the scaffold, suggested by their flattened shape, cell spreading and the fading edges of the cells. The images also demonstrated that cells cling to the material’s fibres through numerous filopodia that were clearly visible within the interfibres spaces (Fig. 12.21). The validity of the scaffold’s design and processing was confirmed by the presence of adhering cells that were restricted to the external rings of the scaffold: cellular adherence was effectively hampered by the occlusion of porosity established at the bottom of the grooves and along the rings wall (Fig. 12.22). Examination of the longitudinal section of the scaffold indicates satisfactory cellular colonisation of the matrix thickness: a remarkable number of cells

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 266

6/6/11 6:50:20 PM

Tracheal tissue regeneration

267

20 µm (a)

2 µm (b)

12.21 ESEM images of the scaffold following a 14-day period of dynamic culture in a bioreactor. Evidence of the good adherence of the cellular component to the scaffold fibres was observed by (a) cell spreading and interfibre colonisation. (b) Magnification of the rectangular white box in (a).

penetrated in depth, so that a considerable number of cell aggregates were observed at approx. 570 mm from the external surface. Immunocytochemical examination of the dynamised scaffold samples, using a monoclonal anticollagen type II antibody, showed the presence of chondrocytes both on the surface and throughout the Degrapol® fibres (Fig. 12.23). The number

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 267

6/6/11 6:50:20 PM

268

Electrospinning for tissue regeneration

20 µm

12.22 ESEM image of the scaffold microstructure at the bottom of the grooves: cellular adherence has been effectively hampered by the occlusion of porosity established during processing.

(a)

(b)

(c)

12.23 Immunocytochemical images showing a clear amount of positive cells on the scaffold surface (white arrows) and their tendency to penetrate deep inside the fibres. The number of collagen type II positive cells was higher with evidence of a thin layer on the external surface (a) and (c) while their numbers decrease thoughout the scaffold thickness (b) (Scale bars = 50 µm).

of collagen type II positive cells was higher, with evidence of a thin layer on the external surface and a decreased number throughout the scaffold thickness. In order to investigate the potential effects of mechanical stimulation on the cells, gene expression for type II collagen, type I collagen, aggrecan and glyceraldehyde 3-phosphate dehydrogenase (GADPH) (control) was determined by a polymerase chain reaction (PCR). Semi-quantitative results of the preliminary tests on gene expression conducted on chondrocytes at

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 268

6/6/11 6:50:20 PM

Tracheal tissue regeneration

269

two different time points, preseeding and post-dynamic culturing in the bioreactor, are reported in Fig. 12.24. The results obtained suggest constant expression of type II collagen in both samples with a marked decrease of type I collagen in the post-dynamic samples. Aggrecan gene expression showed the same aspect in the pre- and post-bioreactor samples. The weak signal in the post-bioreactor samples can reasonably be related to the presence of a poor extracellular matrix produced by the chondrocytes. Data obtained comparing the native cartilage tissue, the pre- and the post-bioreactor samples showed a significant difference between the three tissue types, but only for the expression of type I collagen (p < 0.001). Type II collagen expression was statistically significant between the native and the TET but not between the pre- and post-dynamic culture tissues, which contrasted to the expected results following dynamic stimulation.

12.5.3 In vivo experimental procedures Data obtained after in vivo implantation of the ad hoc designed toothed profile tracheal scaffold in rabbits in an isolated vascular flap using the common carotid artery and the external jugular vein as blood carriers (Fig. 12.25) showed the behaviour of the electrospun polymer in the natural, biological environment with emphasis on degradability, biocompatibility and ability to induce angiogenesis (Brizzola, 2009). The scaffold was made by three 360° rings 2.0 mm in length interposed by 0.5 mm of inter-ring space. The tubular Pre bioreactor

Post bioreactor

GADPH

Collagen II

Collagen I

Aggrecan

12.24 Gene expression for chondrocytes at two different time points: preseeding and post-dynamic culturing in the bioreactor. The presence of GADPH-related frequency bands indicates a satisfactory outcome of the amplification reaction, confirming the expected results for collagen types I and II.

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 269

6/6/11 6:50:20 PM

270

Electrospinning for tissue regeneration

12.25 Vascular axis formed by the common carotid artery (arrowhead) and the internal jugular vein (arrow) of rabbit. The integrity of the vago-sympathetic trunk was carefully maintained.

skeleton measured 6 mm outside and 5 mm inside diameter, respectively, and the entire length was made up by 1 cm of electrospun Degrapol®, with 6/25 mm fibre diameter and 30/80 mm interfibre space and a molecular weight 60.765 Da. A Teflon cylinder was introduced inside the lumen to prevent internal tissue overgrowth. Rabbits were divided in 3 groups: Group 1, Group 2 and Group 3 refers to two, six and eight weeks of implantation, respectively. The tissue invasion started from the vascular axis and proceeded, time depending, towards the anti-pedicle zone. A thin fibroconnective tissue was clearly present around the tracheal scaffold in all animals (Fig. 12.26). Full polymer resorption during the 8 weeks of implantation was not complete. In all samples examined, fibroblasts and matrix in the form of fibrils were in direct contact with the Degrapol® scaffolds (Fig. 12.27(a)–(h)). Initially (Group 1) few cells and fibrillar matrix adhered to the scaffold and spanned within it (Fig. 12.27(a)–(d)). The number and quantity of cells and fibrillar matrix strongly increased with time and after 8 weeks (Group 3) the scaffolds’ external surface was partially concealed (Fig. 12.27(e)) while the polymer fibres were incorporated in a dense fibrous tissue with interposed cells (Fig. 12.27(f), (h)). Histological examination showed a clear presence of cells and extracellular fibrillar matrix, starting from the external side of the scaffolds and moving towards the inner lumen (Fig. 12.28). The cellular invasion proceeded throughout the scaffold thickness (Fig. 12.29(a)–(c)) and visible new blood vessels were evident mostly in the area of maximum tissue deposition (Fig. 12.29(c)). As predicted, the zones of the

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 270

6/6/11 6:50:20 PM

Tracheal tissue regeneration

G1

G1

G2

G2

G3

G3

271

12.26 Macroscopic evaluation of the complex vascular pedicle scaffold for angiogenesis (asterisks) and neo-tissue formation (arrowheads). Rabbits were divided into three groups: G1, G2 and G3, which refers to implantation at two, six and eight weeks, respectively. The images clearly show a progressive time-dependent neovascular tissue growth onto and through the scaffold starting from the vascular axis.

scaffold closest to the vascular carrier were reached by fibroblast-like cells before the scaffold zones located at a greater distance (Acocella, 2007). The pedicle region showed early signs of neo-angiogenesis by means of tissue deposition accompanied by new formation of blood vessels. In all implanted scaffolds, large vascular spaces filled with erythrocytes were observed in close proximity to the vascular pedicle. A few macrophages (inflammatory cells) were involved in the reactive process and were mainly localised in proximity to the blood vessels. There was no evidence of other cells involved in the immune process being present, confirming the good biocompatibility of the polymer and its fibrous structure. Transmission electron microscopy (TEM) evaluation showed cells to be surrounded by collagen fibrils (Fig. 12.29(d) and (e)). The newly synthesised and secreted collagen was organised into tightly packed fibrils

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 271

6/6/11 6:50:21 PM

272

Electrospinning for tissue regeneration

(a)

(b)

(c)

(d)

(e)

(f)

(g)

(h)

12.27 SEM images showing (a)–(d) that the Degrapol® scaffold, after 2 weeks implantation, was partially covered by cells and fibrillar matrix, with cells lying in close contact to the scaffold fibres (arrowheads) and bridging between fibres (arrows) as demonstrated by their visible filopodia. After 8 weeks implantation (e)–(h) cells and fibrillar material have increased in number and quantity and conceal a larger area of the Degrapol® scaffold (Scale bars: (a) 50 µm, (b) 5 µm, (c)–(h), 20 µm).

with a regular and parallel arrangement to form bundles (Fig. 12.29(d) and (e)). Fibroblast-like cells, characterised by a spreading phenotype, penetrated into the surrounding extracellular matrix (Fig. 12.29(e)). As the cells proliferated and migrated centripetally filling the pores of the scaffold, small muscle fibres (Fig. 12.29(f)) and endothelial cells defining capillary structure became visible (Fig. 12.29(f) and (g)). Optical immunofluorescence staining showed a regular and increased deposition of collagen starting from

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 272

6/6/11 6:50:21 PM

Tracheal tissue regeneration

(a)

(b)

273

(c)

12.28 Histological analysis. Scattered early evidence of fibrovascular invasion (a). The presence of large vascular spaces (white asterisks), provides evidence of a neo-angiogenetic process (b,c), filled with a variable number of erythrocytes (black asterisks) (hematoxylin-eosin).

the surface and filling the biopolymer network (Fig. 12.29(h)). At the end of the invasion process the collagen was also localised around the neo-vessels (Fig. 12.29(i)). Optical and electron microscopy, and immunohistochemistry and immunocytochemistry have been used for this purpose. The monoclonal CD31 antibody analysis, detecting progenitor cells (activate endothelial cells-ECs), showed that positive cells were visibly crowded especially on the external surface of the scaffold (Fig. 12.29(j)) while a short time after cell invasion, the expression of CD31 was localised to the endothelial cells forming the blood vessel walls (Fig. 12.29(k)). Cell numbers increased over time, with many of the cells differentiating into muscle fibres, validated by the expression of a-SMA (alpha smooth muscle antibody) (Fig. 12.29(l)). In all samples few inflammatory cells expressing CD14 were visible (Fig. 12.29(m)), while most cells adopting a pronounced migratory and spreading phenotype showed a strong positivity for cathepsin B, a proteolytic enzyme involved in different biological mechanisms, in other words neo-angiogenesis (Fig. 12.29(n)). Our analysis showed that over time there was a migration of CD31 positive cells, which is generally expressed from endothelial progenitor cells derived by an earlier common myeloid progenitor or hematopoietic stem cell. Based on the hypothesis that CD31 is an early indicator during endothelial differentiation, CD31-bright cells were identified as precursor cells moving from the vascular pedicle and then colonising the full thickness of the Degrapol® scaffold. The scaffolds revealed a noticeable amount of neo-tissue formation that progressively substituted the biopolymer fibres by filling the interfibre spaces with cells and extracellular fibrillar matrix made of collagen. Morphological and immunocytochemical characterisation was consistent with endothelial cell migration, spatial disposition and differentiation of migrating stem cells in various cell types. The positivity of cathepsin B in the present cells advocates its properties in promoting angiogenesis through a process of migration and invasion, already described by others in both in vitro and in vivo studies (Lutgens, 2007; Premzl, 2006). Cathepsin B seems to be implicated in different pathological

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 273

6/6/11 6:50:21 PM

274

Electrospinning for tissue regeneration

(a)

(c)

(b)

(d)

(e)

(f)

(g)

(i)

(j)

(k)

V

(h)

(o)

(l)

(m)

(n)

(p)

12.29 Semi-thin cross-section of excised tracheal scaffold (a)–(c). Progressively, cells and fibrillar material invade centripetally the Degrapol® scaffold. Transmission electron microscope images showing a detail of the Degrapol® scaffold covered by cells and fibrillar material. A large number of collagen bundles longitudinally sectioned (d), with fibroblast-like cells (e) surrounded by collagen bundles cross-sectioned (arrowheads), small muscle fibres (f) and minute blood vessels (g) were visible. Prosthesis crosssectioned (h, i) showing an increased deposition of collagen filling the biopolymer network (h) and surrounding the neovessels (i, v). Immunofluorescence (j, k) with monoclonal anti-CD 31, antiSMA (l), anti CD14 (m), and anti-cathepsin B (n) muscle fibres (l), macrophages (m) and cells characterised by cathepsin B production (n) are visible (white arrowheads). Cryosections of Degrapol® scaffold before (o) and after (p) incubation with cathepsin B enzyme for 1 h at 37°C. The fibres of biopolymer were broken and reduced in dimension (p) owing to enzyme degradation. (Scale bars: (a)–(c), (i)–(p) 100 µm, (d)–(f) 2 µm, (g) 4 µm, (h) 50 µm).

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 274

6/6/11 6:50:21 PM

Tracheal tissue regeneration

275

processes such as tumour angiogenesis, neurodegeneration or abdominal aorta aneurysm (Abisi, 2007) but how it regulates angiogenesis, cell proliferation, invasion and apoptosis is still poorly understood. Moreover, our simple enzymatic activity analysis showed that cathepsin B could be responsible for the polymer degradation similar to degradation of the biological matrix. Our data showed the expression of cathepsin B in progenitor cells and expression of CD31 and SMA, which is consistent with a migratory and angiogenetic cellular phase. To our knowledge, it is the first time that cathepsin B has been found during an in vivo evaluation of a polymer scaffold undergoing degradation (Santerre, 2005). Further studies on the interaction between cathepsin B and Degrapol®, in terms of chemical structure, would improve our knowledge of which part of the polymer is subjected to the degradation process.

12.6

Conclusions

A new tubular tracheal construct, basing on the anatomy of the native trachea, has been designed, using the electrospinning technique. This method allowed us to realise a microstructured scaffold with an optimal geometry for cell adhesion, growth and proliferation. The toothed profile exhibited theoretical functional characteristics that prevent kinking or collapse during eventual physiological movements of the implanted scaffold. The electrospinning process resulted in a 3D microfibrillar network with an optimal fibre diameter (6/25 mm), an interfibre space of 30/80 mm and open porosity on the surface defined as the gold standard in allowing appropriate gas flow and exchange of growth factors and cell nutrients (Sanders, 2005). Furthermore a very thin microporous inner layer was added to the scaffold to prevent tissue ingrowth and the subsequent reduction in the scaffold’s lumen. The absence of a foreign body reaction after implantation in animals confirmed that the architecture of the structure was similar to a natural extracellular matrix, indicating that the electrospun scaffold was suitable as a tissue substitute. This tracheal engineered scaffold showed a specific profile with open porosity (seeding area) on the rings’ surface and closed porosity over the intra-annular walls, controlling and guiding the cellular attachment and proliferation, which then allowed the residing cells to build up their own ECM as the polymer fibres were hydrolysed and degraded over time. The very thin microporous inner layer proved to be very useful in enabling gas flow and exchange of growth factors and cellular nutrients, whilst at the same time preventing tissue ingrowth and thus avoiding reduction of the scaffold’s lumen. The use of a dynamic culture bioreactor system documented a very high efficiency in the cell-seeding procedure. Scaffold movement did not hamper cell adhesion to the fibre surface and seemed to promote cell migration

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 275

6/6/11 6:50:21 PM

276

Electrospinning for tissue regeneration

throughout the thickness of the scaffold fibres, which favourably supported cell phenotype. The results of the in vivo studies demonstrated that this bioresorbable polymer provided a good substrate for fibrous tissue deposition and formation of neo-angiogenesis throughout the tubular thickness of the scaffold. This is essential for scaffold stability and epithelial growth. Looking forward it is possible to theorise about a new scaffold design that will take advantage from the host biological process to guide more sophisticated and physiological tissue regeneration.

12.7

Acknowledgements

The authors thank abmedica s.p.a., Lainate, Italy for providing Degrapol®.

12.8

References

Abisi S, Burnand KG, Waltham M, Humphries J, Taylor PR and Smith A (2007), ‘Cysteine protease activity in the wall of abdominal aortic aneurysms’, J Vasc Surg, 46(6), 1260–6. Acocella F, Brizzola S, Valtolina C, Scanziani E, Marchesi F, Mantero S, Garreau H and Vert M (2007), ‘Prefabricated tracheal prosthesis with partial biodegradable materials: a surgical and tissue engineering evaluation in vivo’, J Biomater Sci Polym Ed, 18(5), 579–94. Agrawal CM and Ray RB (2001), ‘Biodegradable polymeric scaffolds for musculoskeletal tissue engineering’, J Biomed Mater Res, 55, 141–50. Allemann F, Mizuno S, Eid K, Yates KE, Zaleske D and Glowacki J (2001), ‘Effects of hyaluronan on engineered articular cartilage extracellular matrix gene expression in 3-dimensional collagen scaffolds’, J Biomed Mater Res, 55, 13–19. Begis D, Delpuech C, Le Tallec P, Loth L, Thiriet M and Vidrascu M (1988), ‘A finiteelement model of tracheal collapse’, J Appl Physiol, 64(4), 1359–68. Belsey R (1950), ‘Resection and reconstruction of the intrathoracic trachea’, Br J Surg, 38, 200–5. Benya PD and Shaffer JD (1982), ‘Dedifferentiated chondrocytes reexpress the differentiated collagen phenotype when cultured in agarose gels’, Cell, 30, 215–24. Bonaventure J, Kadhom N, Cohen-Solal L, Ng KH, Bourguignon J, Lasselin C and Freisinger P (1994), ‘Re-expression of cartilage-specific genes by dedifferentiated human articular chondrocytes cultured in alginate beads’, Exp Cell Res, 212, 97–104. Borkenhagen M,. Stoll RC, Neuenschwander P, Suter UW and Aebischer P (1998), ‘In vivo performance of a new biodegradable polyester urethane system used as nerve guidance channel’, Biomaterials, 19(23), 2155–65. Brizzola S, de Eguileor M, Brevini T, Grimaldi A, Congiu T, Neuenschwander P and Acocella F (2009), ‘Morphologic features of biocompatibility and neoangiogenesis onto a biodegradable tracheal prosthesis in an animal model’, Interact Cardiovasc Thorac Surg, 8(6), 610–4. Cohen RC, Filler RM, Konuma K, Bahoric A, Kent G and Smith C (1985), ‘The successful reconstruction of thoracic tracheal defects with free periosteal grafts’, J Pediatr Surg, 20, 852–8.

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 276

6/6/11 6:50:21 PM

Tracheal tissue regeneration

277

Dahlberg L and Kreicbergs A (1991), ‘Demineralized allogenic bone matrix for cartilage repair’, J Orthop Res, 9, 11–19. Danielsson C, Ruault S, Simonet M, Neuenschwander P and Frey P (2006), ‘Polyesterurethane foam scaffold for smooth muscle cell tissue engineering’, Biomaterials, 27(8), 1410–15. Davilla JC, Lautsch EV and Palmer TE (1968), ‘Some physical factors affecting the acceptance of synthetic materials as tissue implants’, Ann NYAS, 146, 138–47. Doshi J and Reneker DH (1995), ‘Electrospinning process and applications of electrospun fibers’, J Electrostat, 35, 151–60. Fonkalsrud EW and Sumida S (1971),‘Tracheal replacement with autologous esophagus for tracheal stricture’, Arch Surg, 102, 139–42. Grillo HC (1965), ‘Circumferential resection and reconstruction of the mediastinal and cervical trachea’, Ann Surg, 162, 374–88. Grillo HC (2002), ‘Tracheal replacement: a critical review’, Ann Thorac Surg, 73(6), 1995–2004. Grillo HC (2003), ‘Development of tracheal surgery: a historical review. Part 2: Treatment of tracheal diseases’, Ann Thorac Surg, 75(3), 1039–47. Har-El G, Krespi YP and Goldsher M (1989), ‘The combined use of muscle flaps and alloplasts for tracheal reconstruction’, Arch Otolaryngol Head Neck Surg, 115, 1310–13. Hohman MM, Shin M, Rutledge GC and Brenner MP (2001), ‘Electrospinning and electrically forced jets. II. Applications’, Phys Fluids, 13, 2221–36. Jackson TL, O’Brien EJ, Tuttle W and Meyer J (1950), ‘The experimental use of homogenous tracheal transplants in the restoration of continuity of the tracheobronchial tree’, J Thorac Surg, 20, 598–612. Jacobs JR (1988), ‘Investigations into tracheal prosthetic reconstruction’, Laryngoscope, 98, 1239–45. Jakob M, Démarteau O, Schäfer D, Hintermann B, Dick W, Heberer M and Martin I (2001), ‘Specific growth factors during the expansion and redifferentiation of adult human articular chondrocytes enhance chondrogenesis and cartilaginous tissue formation in vitro’, J. Cell. Biochem, 81, 368–77. Jansen JA, von Recum AF, van der Waerden JP and de Groot K (1992), ‘Soft tissue response to different types of sintered metal fibreweb materials’, Biomaterials, 13, 959–68. Kojima K, Bonassar LJ, Ignotz RA, Syed K, Cortiella J and Vacanti CA (2003a), ‘Comparison of tracheal and nasal chondrocytes for tissue engineering of the trachea’, Ann Thorac Surg, 76, 1884–8. Kojima K, Bonassar LJ, Roy AK, Mizuno H, Cortiella J and Vacanti CA (2003b), ‘A composite tissue-engineered trachea using sheep nasal chondrocytes and ephithelial cells’, FASEB J, 17, 823–8. Kojima K, Bonassar LJ, Roy AK, Vacanti CA and Cortiella J (2002), ‘Autologous tissue-engineered trachea with sheep nasal chondrocytes’, J Thorac Cardiovasc Surg, 123, 1177–84. Kon M and van den Hooff A (1983), ‘Cartilage tube formation by perichondrium: a new concept for tracheal reconstruction’, Plast Reconstr Surg, 72, 791–7. Lambert RK, Baile EM, Moreno R, Bert J and Paré PD (1991), ‘A method for estimating the Young’s modulus of complete tracheal cartilage rings’, J Appl Physiol, 70(3), 1152–9. Letang E, Sanchez-Lloret J, Gimferrer JM, Ramirez J and Vicens A (1990), ‘Experimental

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 277

6/6/11 6:50:21 PM

278

Electrospinning for tissue regeneration

reconstruction of the canine trachea with a free revascularized small bowel graft’, Ann Thorac Surg, 49, 955–8. Lutgens SP, Cleutjens KB, Daemen MJ and Heeneman S (2007), ‘Cathepsin cysteine proteases in cardiovascular disease’, FASEB J, 21(12), 3029–41. (Review). Macchiarini P, Jungebluth P, Go T, Asnaghi MA, Rees LE, Cogan TA, Dodson A, Martorell J, Bellini S, Parnigotto PP, Dickinson SC, Hollander AP, Mantero S, Conconi MT and Birchall MA (2008), ‘Clinical transplantation of a tissue-engineered airway’, Lancet, 372, 2023–30. Okumus A, Cizmeci O, Kabakas F, Kuvat SV, Bilir A and Aydin A (2005), ‘Circumferential trachea reconstruction with a prefabricated axial biosynthetic flap: experimental study’, Int J Pediatr Otorhinolarynol, 69, 335–44. Osada H, Takeuchi S, Kojima K and Yamate N (1994), ‘The first step of experimental study on hybrid trachea: use of cultured fibroblasts with artificial matrix’, J Cardiovasc Surg, 35(Suppl. 1–6), 165–8. Paquay YC, de Ruijter JE, van der Waerden JP and Jansen JA (1996), ‘A one stage versus two stage surgical technique: Tissue reaction to a percutaneous device provided with titanium fiber mesh applicable for peritoneal dialysis’, ASAIO J, 42, 961–7. Ponticiello MS, Schinagl RM, Kadiyala S and Barry FP (2000), ‘Gelatin based resorbable sponge as a carrier matrix for human mesenchymal stem cells in cartilage regeneration therapy’, J Biomed Mater Res, 52, 246–55. Premzl A, Turk V and Kos J (2006), ‘Intracellular proteolytic activity of cathepsin B is associated with capillary-like tube formation by endothelial cells in vitro’, J Cell Biochem, 97(6), 1230–40. Raimondi MT, Boschetti F, Falcone L, Migliavacca F, Remuzzi A and Dubini G (2004), ‘The effect of media perfusion on three-dimensional cultures of human chondrocytes: integration of experimental and computational approaches’, Biorheology, 41(3), 401–10. Rains JK, Bert JL, Roberts CR and Paré PD (1992), ‘Mechanical properties of human tracheal cartilage’, J Appl Physiol, 72, 219–25. Riboldi SA, Sampaolesi M, Neuenschwander P, Cossu G and Mantero S (2005), ‘Electrospun degradable polyesterurethane membranes: potential scaffolds for skeletal muscle tissue engineering’, Biomaterials, 26(22), 4606–15. Roberts CR and Parè PD (1991), ‘Composition changes in human tracheal cartilage in growth and aging, including changes in proteoglycan structure’, Am J Physiol, 261(2 Pt 1), L92–101. Ruszymah BHI, Chua K, Latif MA, Hussein FN and Saim AB (2005), ‘Formation of in vivo tissue engineered human hyaline cartilage in the shape of a trachea with internal support’, Int J Pediatr Otholaryngol, 69, 1489–95. Saad B, Ciardelli G, Matter S, Welti M, Uhlschmid GK, Neuenschwander P and Suter UW (1996), ‘Characterization of the cell response of cultured macrophages and fibroblasts to particles of short-chain poly[(R)-3-hydroxybutyric acid]’, J Biomed Mater Res, 30(4), 429–39. Saad B, Hirt TD, Welti M, Uhlschmid GK, Neuenschwander P and Suter UW (1997), ‘Development of degradable polyesterurethanes for medical applications: in vitro and in vivo evaluations’, J Biomed Mater Res, 36(1), 65–74. Saad B, Moro M, Tun-Kyi A, Welti M, Schmutz P, Uhlschmid GK, Neuenschwander P and Suter UW (1999), ‘Chondrocyte-biocompatibility of degrapol-foam: in vitro evaluations’, J Biomater Sci Polym Edn, 10(11), 1107–19. Sanders JE, Stiles CE and Hayes CL (2000), ‘Tissue response to single-polymer fibers

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 278

6/6/11 6:50:22 PM

Tracheal tissue regeneration

279

of varying diameters: evaluation of fibrous encapsulation and macrophage density’, J Biomed Mater Res, 52, 231–7. Sanders JE, Lamont SE, Mitchell SB and Malcolm SG (2005), ‘Small fiber diameter fibro-porous meshes: tissue response sensitivity to fiber spacing’, J Biomed Mater Res, 72A, 335–42. Santerre JP, Woodhouse K, Laroche G and Labow RS (2005), ‘Understanding the biodegradation of polyurethanes: from classical implants to tissue engineering materials’, Biomaterials, 26(35), 7457–70. Scherer MA, Ascherl R, Geissdörfer K, Mang W, Blümel G, Lichti H and Fraefel W (1986), ‘Experimental biosynthetic reconstruction of the trachea’, Arch Otorhinolarynol, 243, 215–23. Schimdt MB, Mow VC, Chun LE and Eyre DR (1990), ‘Effects of proteoglycan extraction on the tensile behaviour of articular cartilage’, J Orthop Res, 8, 353–63. Stankus JJ, Guan J and Wagner WR (2004), ‘Fabrication of biodegradable elastomeric scaffolds with sub-micron morphologies’, J Biomed Mater Res A, 70(4), 603–14. Thomson RC, Yaszemski MJ and Mikos AG (1997), ‘Polymer scaffold processing’, in Principles of Tissue Engineering, Lanza RP, Langer R and Chick WL (eds), R.G. Landes, Austin, TX. Vacanti CA, Paige KT, Kim WS, Sakata J, Upton J and Vacanti JP (1994), ‘Experimental tracheal replacement using tissue-engineered cartilage’, J Pediatr Surg, 29(2), 201–4, discussion 204–5. Vunjak-Novakovic G, Martin I, Obradovic B, Treppo S, Grodzinsky AJ, Langer R and Freed LE (1999), ‘Bioreactor cultivation conditions modulate the composition and mechanical properties of tissue-engineered cartilage’, J Orthop Res, 17, 130–8. Watt FM and Dudhia J (1988), ‘Prolonged expression of differentiate phenotype by chondrocytes cultured at low density on a composite substrate of collagen and agarose that restricts cell spreading’, Differentiation, 38, 140–7. Yang L, Korom S, Welti M, Hoerstrup SP, Zünd G, Jung FJ, Neuenschwander P and Weder W (2003), ‘Tissue engineered cartilage generated from human trachea using DegraPol scaffold’, Eur J Cardiothorac Surg, 24(2), 201–7.

© Woodhead Publishing Limited, 2011

Electrospinning-Bosworth-12.indd 279

6/6/11 6:50:22 PM