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  Bioprinting of arti 󿬁 cial blood vessels: current approachestowards a demanding goal Eva Hoch a  , Günter E.M. Tovar  a,b, * and Kirsten Borchers b, * a Institute for Interfacial Process Engineering and Plasma Technology IGVP, University of Stuttgart, Stuttgart, Germany b Fraunhofer Institute for Interfacial Engineering and Biotechnology IGB, Stuttgart, Germany* Corresponding author. Fraunhofer Institute for Interfacial Engineering and Biotechnology IGB, Nobelstr. 12, 70569 Stuttgart, Germany. Tel: +49-711-9704121;fax: +49-711-9704200; e-mail: kirsten.borchers@igb.fraunhofer.de (K. Borchers); Institute for Interfacial Process Engineering and Plasma Technology IGVP, Universityof Stuttgart, Nobelstr. 12, 70569 Stuttgart, Germany . Tel: +49-711-9704109 ; fax: +49-711-9704200; e-mail: guenter.tovar@igvp.uni-stuttgart.de (G. Tovar).Received 6 May 2014; accepted 8 May 2014 Summary  Free-form fabrication techniques, often referred to as  ‘ 3D printing ’ , are currently tested with regard to the processing of biological and bio-compatible materials in general and for fabrication of vessel-like structures in particular. Such computer-controlled methods assemble 3Dobjects by layer-wise deposition or layer-wise cross-linking of materials. They use, for example, nozzle-based deposition of hydrogels andcells, drop-on-demand inkjet-printing of cell suspensions with subsequent cross-linking, layer-by-layer cross-linking of synthetic or bio-logical polymers by selective irradiation with light and even laser-induced deposition of single cells. The need of vessel-like structures hasbecome increasingly crucial for the supply of encapsulated cells for 3D tissue engineering, or even with regard to future application suchas vascular grafts. The anticipated potential of providing tubes with tailored branching geometries made of biocompatible or biologicalmaterials pushes future visions of patient-speci 󿬁 c vascularized tissue substitutions, tissue-engineered blood vessels and bio-based vasculargrafts. We review here the early attempts of bringing together innovative free-form manufacturing processes with bio-based and bio-degradable materials. The presented studies provide many important proofs of concepts such as the possibility to integrate viable cellsinto computer-controlled processes and the feasibility of supplying cells in a hydrogel matrix by generation of a network of perfused chan-nels. Several impressive results in the generation of complex shapes and high-aspect-ratio tubular structures demonstrate the potential of additive assembly methods. Yet, it also becomes obvious that there remain major challenges to simultaneously match all material require-ments in terms of biological functions (cell function supporting properties), physicochemical functions (mechanical properties of theprinted material) and process-related (viscosity, cross-linkability) functions, towards the demanding goal of biofabricating arti 󿬁 cial bloodvessels. Keywords:  Bioprinting  ã  Tissue engineering  ã  Arti 󿬁 cial blood vessels INTRODUCTION Tissue engineering (TE) aims for the  in vitro  generation of func-tional, arti 󿬁 cial tissues and organs, which may serve as implants in vivo  [1] or as  in vitro  test systems [2 – 4]. In the classical approach,cells are seeded onto 3D matrices [1] that are either bio-basedsuch as decellularized donor tissue [5 – 8] or synthetic materials fab-ricated byelectrospinning, freeze drying, foaming or rapid prototyp-ing technologies [9 – 12]. Another strategy is the encapsulation of cells into 3D hydrogels [13 – 15].Based on these techniques, simpleand thin tissue mimics of, for example, skin [16,17],cartilage [4,18, 19] or cornea [20] have been realized, partially with the help of  special perfusing bioreactors [21, 22]. Yet, the size of engineered tissues is restricted due to the inability to incorporate a suf  󿬁 cientblood vessel system for nutrient and oxygen supply to the cells[23].The native vascular system is a complex network of bloodvessels of various sizes. The diameters of blood vessels range fromthe centimetre to the micrometre scale, i.e. from  2.5 cm to  20 μ m for the aorta to very  󿬁 ne capillaries, respectively [24].The inner face of the whole vascular system is lined with a mono-layer of   󿬂 at cells, the endothelial cells (ECs). This EC layer preventsthrombogenesis in the blood  󿬂 ow and represents a cellular barrierthat controls the exchange of molecules from blood and tissue.Branching of vessels occurs consistent with certain rules that guar-antee haemostasis, for example, perpetuation of a homogeneouswall shear stress [25]. The maximum distance between two capil-laries is 200 µm, which correlates to the diffusion limitation of oxygen [23].With a view to the generation of larger arti 󿬁 cial tissue and im-plantation of tissue into patients, current research focuses on the in vitro  generation of functional, arti 󿬁 cial blood vessel systems.Additionally, the lack of small-diameter vascular grafts motivatesthe search for technologies and materials to provide biomimeticvascular structures for new generative therapies.As one possible approach, cell-based strategies focus on thepromotion of angiogenesis by (self-)assembly of vascular cells. Forexample Moya  et al . have recently demonstrated the assembly of ECs to form an interconnected and perfusable human capillarynetwork as part of a micro 󿬂 uidic device as they adjusted adequate © The Author 2014. Published by Oxford University Press on behalf of the European Association for Cardio-Thoracic Surgery. All rights reserved.     R    E    V    I    E    W European Journal of Cardio-Thoracic Surgery 46 (2014) 767 – 778  REVIEW doi:10.1093/ejcts/ezu242 Advance Access publication 26 June 2014  pressure drops between  ‘ arterial ’  and  ‘ venular ’  channels [26]. Theircombination of de 󿬁 ned inlet and outlet branches as provided bythe micro 󿬂 uidic device with the formation of a genuine capillarynetwork within the device enabled perfused vascular networks  invitro , but in its current form is not useful for the assembly of 3Dtissue or implantation into living organisms. With respect to  invivo  application of arti 󿬁 cial tissue, it has a been observed in earlystudies by Tremblay  et al . [27] and has also been acknowledged in recent studies performed, for example, by Nunes  et al . [28] that tissue structures containing fragments of capillaries connectedmuch faster to the native vascular system of a host upon implant-ation than tissue without such vascular prestructures. Yet, suchconstructs, on the other hand, do not provide continuous tubularvessel systems  in vitro  and, therefore, they cannot be used for thesupply of surrounding cells as needed for 3D TE.Thus, the essential requirements for fabrication of perfusablevascularized tissue seem to be both, to have de 󿬁 ned inlet andoutlet branches and to provide a capillary-like network with bio-mimetic vessel geometry for undisturbed  󿬂 uid dynamics [29].Therefore, additional approaches provide ready-made tubes ortubular networks as a 3D matrix for supplying cells in 3D culture.These may be either based on the reuse of biologically derived,decellularized vessel systems [30] or synthetically manufacturedtubular scaffolds [31]. To develop arti 󿬁 cial structures that performas well as natural ones, we need fabrication processes that do notset any limits to the generation of structures and shapes, andmaterials that allow for tailoring of their physical, chemical andbiological properties. Bioprinting technology, which applies free-form fabrication methods to deposit scaffold materials and cells toform digitally de 󿬁 ned 3D structures [32], constitutes a basicallynovel approach to approach this so far unsolved issue: There arevarious types of 3D manufacturing techniques addressing a broadrange of structure sizes and several have already been used forbiofabrication purposes as well. THE POTENTIALOF BIOPRINTING TECHNIQUES Bioprinting means the generation of bioengineered structures byadditive manufacturing of biological and biologically relevantmaterials with the use of computer-aided transfer and build-upprocesses [33 – 35]. A variety of solid freeform fabrication techni-ques are available. They can be divided into laser-based methods,printer-based methods and nozzle-based methods.Nozzle-based systems use pressure-assisted syringes to depositcontinuous strands of materials. They typically address structuresizes in the centimetre range with resolutions of several hundredmicrons [36]. Printer-based systems include thermal and piezo-electric inkjet printing. These drop-on-demand systems generatesmall droplets of low viscosity  ‘ bio ’  ink [37]. Typical resolutions arein the range of    85 – 300 µm and can be used for generation of millimetre to centimetre constructs [35]. Laser-based systems in-cluding stereolithography and multiphoton polymerization as wellas digital light processing (DLP), a method based on digital micro-mirror devices [38], use light for site-selective curing of photo-sensitive prepolymers in a bath with resolutions down to thesub-micron range [39].Three parallel bioprinting-based approaches are followedtowards the aim of the generation of arti 󿬁 cial blood vessels: (i)generation of bulk matrices with integrated channels as perfusablematrices, (ii) cell patterning into line structures for self-assembly of interconnected vessel systems and (iii) generation of free-standingtubular structures, with and without encapsulated cells, whichmay serve as arti 󿬁 cial blood vessels (Table 1). PERFUSABLE MATRICES The most straightforward approach to perfusable tissue mightbe the generation of a network of interconnected channels withinthe tissue matrix. Such channels may be used as supply system forcells within the surrounding matrix and may additionally beseeded with ECs. Early works used moulds for preparing sacri 󿬁 cialstructures in order to fabricate micro 󿬂 uidic networks, which thenallowed the transport of macromolecules into surrounding hydro-gels under low driving pressure differences. Such studies resultedin zones of high cell viability up to 200 µm away from the perfusedchannels and thus demonstrate the effectiveness of the strategy.Yet, the soft lithographic techniques that were applied for channelgeneration require multiple processing steps and are limited toplanar networks and stacks of planar networks [65 – 67]. Anotherapproach presented by Bellan  et al . used melt-spun  󿬁 bres as a sac-ri 󿬁 cial mould and thus achieved a 3D channel network, yet inrandom spatial orientation [40].For rapid casting of a de 󿬁 ned pattern of vascular channels,Miller  et al . developed a sacri 󿬁 cial material based on carbohy-drates, which they printed in  󿬁 laments using a syringe and aRepRap Mendel 3D printer [41]. They used mixtures out of glucose, sucrose and dextran (86 kDa) as  ‘ liquid glass ’  and pro-duced self-supporting perpendicular lattices as well as curved  󿬁 la-ments and Y-junctions with  󿬁 bre diameters in the range of 150 – 750 µm (Fig. 1). Such glass lattices were coated with thin layers of poly(lactid- co -glycolid) (PLGA) and subsequently encapsulated inthe presence of living cells using a wide range of natural and syn-thetic matrix materials such as  󿬁 brin, agarose and Matrigel®, andpoly(ethylene glycol) (PEG)-based materials. After cross-linking of the matrix, the carbohydrate support was dissolved in aqueousmedia. Thereby, the PLGA layer prevented carbohydrate solutionfrom  󿬂 owing through the cell-laden hydrogel. The resulting chan-nels allowed non-leaking perfusion of human blood. ECs wereseeded into the channels and lined the walls of the completenetwork. The authors observed sprouting of new capillaries fromthe channel walls into the surrounding  󿬁 brin gels which had beenladen with 10T1/2 cells. In agarose gels and even in gels madefrom synthetic PEG hydrogels, which were modi 󿬁 ed with the RGDcell recognition peptide, various cell types were live and active atthe gel perimeter and approximately 1 mm around the perfusionchannels. For comparison, within gels without channels, cells wereonly active at the gel slab perimeter but not at the gel core.Lee  et al . dispensed heated gelatin solution for the generationof sacri 󿬁 cial elements and ice-cooled collagen hydrogel precursorat a pH of 4.5 to produce collagen layers, which were then cross-linked by spraying NaHCO 3  solution on top [43]. Dermal  󿬁 bro-blasts grown in such collagen scaffolds containing  󿬂 uidic channelsalso showed signi 󿬁 cantly elevated cell viability compared with theones without any channels, proving again the effectiveness of in-tegrating vascular structures for 3D tissue reconstruction.In their approach to fabricate a free-form 3D channel network within a hydrogel, Wu  et al . used a reservoir of photo-cross-linkable, acrylated Pluronic F127® (PF127, 20 – 25% w/w) anddeposited sacri 󿬁 cial  󿬁 lament networks from nonacrylated PF127directly into such reservoirs [42]. PL127 is an interesting materialfor manufacturing purposes because it shows unusual properties,i.e., it is liquid at 4°C and solidi 󿬁 es when warmed. It is one of the E. Hoch  et al . / European Journal of Cardio-Thoracic Surgery 768  Table 1:  Overview over biofabrication methods and materials used in present approaches for the generation of vascular structures Bioprinting technique Building material/bulk hydrogel Sacrificial material/supportivematerialCell type Channel diameter Ref.Perfusable matricesNozzle-based:Modified cotton candy machineEnzymatically cross-linked gelatin Shellac  –  Diameter: 17 ±19 µm [40]Nozzle-based:RepRap Mendel 3D printerPEG diacrylate+acrylate-PEG-RGDS; Matrigel®;agarose; alginate; fibrinCarbohydrate glass(glucose +sucrose+dextran)HUVEC, 10T1/2 cells Diameters:150 – 750 µm, straightand curved channels[41]Nozzle-based Acrylated Pluronic F127® Pluronic F127®  –  18 – 1200 µm [42]Nozzle-based Collagen Gelatin Primary HDFs µm to mm range [43]Nozzle-based Mixtures of gelatin, alginate, chitosan, fibrinogen, HA  –  Rat primary hepatocytes,ADSCsmm range [44]Cell patterning for autonomous vessel formationLaser-based:BioLP – –  HUVEC, HUVSMC µm range (single-cellpositioning)[45]Laser-based:BioLPAlginate, Matrigel®, fibrin  –  Rabbit carcinoma cell lineB16, HUVEC cell lineEathy926µm to mm range [46]Drop-on-demand:HP Deskjet 500Fibrin  –  HMVEC µm range [47]Nozzle-based:BioAssembly ToolCollagen  –  RFMF mm range [48]Free-standing tubular structuresLaser-based:Two-photon polymerization α , ω -Polytetrahydrofuranether-diacrylate  – –  Diameter: 10 – 100 µm [49]Laser-based:Digital light processingFormulations based on urethan-diarylate, hydroxylethylacrylate as reactive diluent, ethylene glycolbisthioglycolate as the chain transfer agent –  Diameter: 4 µmWall: 2 mmHeight: 2 mm[38]Drop-on-demand:Thermal inkjet, SEAjet ™ Alginate Ca solution +HA/PVA  –  Diameter:50 – 1000 µm[50, 51] Drop-on-demand:Thermal inkjet, SEAjet ™ Alginate Ca solution +PVA HeLa Diameter: 200 µm [52]Drop-on-demand:Thermal inkjet, HP697cAlginate Alginate solution Rat SMC Diameter: 2 mmWall: 2 mmHeight: 2 mm[53]Drop-on-demand:Piezoelectric MicroFabMJ-ABL-01-120-6MX dispenseheadAlginate Ca solution NIH 3T3 fibroblasts [54]Drop-on-demand:Custom-made, syringe+switchablevalve applied droplet volumeof 116 nlAgarose Hydrophobic high-densityperfluorotributylamine C 12 F 27 NHuman MG-63 andhuman MSCDiameter:  7 mmWall:  1.5 mmHeight:  4 cm[55]Nozzle-based:Novogen MMX Bioprinter,OrganovoNovoGel  –  Mouse embryonicfibroblasts[56]Nozzle-based:Fab@HomeThiol-modified HA+thiol-modified gelatin with goldnanoparticles or tetra-acrylated PEG as cross-linkerHA NIH 3T3 fibroblasts Diameter: 3 – 5 mmWall: 1 – 2 mmHeight: 1 – 2 cm[57]Nozzle-based:Fab@HomeMethacrylated HA+methacrylated gelatin HA Human hepatoma cellsHepG2 C3ADiameter: 1 – 2 mmWall: 2 mm[58] Continued  REVIEW E   .H o c h    e  t     al        . /   E   ur   o  p e a n  J    o ur  n a l     of     C  a r   d  i     o-T  h   or   a  c i     c  S   ur    g er    y  7   6   9    few synthetic polymer materials that are approved for clinicalapplications by the American Food and Drug Association FDA[68]. However, the presented study does not discuss the integra-tion of cells into the constructs, and other studies revealed non-cytotoxicity of PF127 gels only up to 20%. [69, 70] Therefore, we consider the use of acrylated PF 127 as a cross-linkable matrix ma-terial to provide a proof of concept of omnidirectional  󿬁 lamentextrusion within a hydrogel precursor matrix rather than a realmatrix for 3D TE.Li  et al . introduced bio-based material systems that were suitedfor vertical assembly of hollow channels without the use of a sup-portive sacri 󿬁 cial material (Fig. 2). The authors used different com-binations of mixed gelatin/alginate/chitosan/  󿬁 brinogen hydrogelsas building materials. Their printed structures were  󿬁 rst stabilizedby the sol – gel transition of the gelatin as the mixtures (37°C) wereextruded from the syringe needle into a low-temperature fabrica-tion chamber (6 – 8°C). The alginate, chitosan and  󿬁 brinogen werethen subsequently cross-linked using thrombin, CaCl 2,  Na 5 P 3 O 10 and glutardialdehyde [44]. CELL PATTERNING FOR SELF-ASSEMBLYOFINTERCONNECTED VESSEL SYSTEMS The second approach to generate blood vessels in arti 󿬁 cial tissuedeals with the ability of ECs to organize into blood vessels autono-mously (angiogenesis).  In vivo  angiogenesis occurs as a responseto hypoxia, for example, in growing tissue when ECs sprout fromexisting blood vessels to form new capillaries by the curling of asingle or few cells [71]. While for  in vitro  TE, an interconnectedsupply system has to be instantaneously available, genuine vascu-lar structures may mature in parallel for the fast integration of theengineered tissue to the host tissue upon implantation. It hasbeen observed that cultivation of ECs that were randomly distribu-ted within 3D, growth factor-rich hydrogels, e.g. Matrigel®, led toformation of an undirected, polygonal cellular network [72]. Thus,it is expected that prestructuring of vascular cells may guide thedirection of lumen formation and growth into an interconnectedcapillary system. By the application of digital patterning techni-ques, directed cell deposition can be achieved and branchingstructures with biomimetic branching geometries may be designed.Towards this aim, biological laser printing (BioLP), also referredto as laser-assisted bioprinting (LAB), allows for deposition of indi-vidual cells. This method is based on laser-induced forward trans-fer of biological materials and cells. Wu and Ringeisen [45] andGuillotin  et al . [46] used LAB for positioning of individual human umbilical vein endothelial cells (HUVECs). Wu and Ringeisendesigned de 󿬁 ned branch and stem structures and observed thatdeposited HUVECs connected with each other within 1 day (Fig. 3).However, the connected branches built from HUVECs alone didnot last more than a few days. Stabilization of the cellular network structure for at least 9 days occurred only with the deposition of anadditional cell layer of human umbilical vein smooth muscle cells(HUVSMCs) on top of HUVECs. This con 󿬁 rms similar results fromother labs and indicates that assembly of ECs with additional vascu-larcells, such as pericytes and smooth muscle cells, is crucial for thegeneration of stable, functional arti 󿬁 cial blood vessels and function-al tissue substitutes [73,74]. While laser-based technologies allow for precise positioning of individual cells and thereby the generation of structures with reso-lution at the cellular level, technologies such as inkjet printing or        T     a       b       l     e       1     :     C   o   n    t    i   n   u   e    d 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     –     –     D    i   a   m   e    t   e   r   :    2      –     4   m   m    W   a    l    l   :    1   m   m    H   e    i   g    h    t   :    2    0   m   m    T   o    t   a    l   s    i   z   e   :    6    7   ×    4    2   ×    8   m   m     ³     [    6    0     ]    N   o   z   z    l   e  -    b   a   s   e    d   :    B    i   o    S   c   a    f    f   o    l    d   e   r    3    D    F   s   y   s    t   e   m    S    Y    S   +    E    N    G ,    G   e   r   m   a   n   y    M   e    t    h   a   c   r   y    l   a    t   e    d   g   e    l   a    t    i   n    /   g   e    l    l   a   n   g   u   m    h   y    d   r   o   g   e    l      –     –     D    i   a   m   e    t   e   r   :    2      –     4   m   m    W   a    l    l   :    1   m   m    H   e    i   g    h    t   :    2    0   m   m    T   o    t   a    l   s    i   z   e   :    6    7   ×    4    2   ×    8   m   m     ³     [    6    0     ]    N   o   z   z    l   e  -    b   a   s   e    d   :    M   u    l    t    i   c   e    l    l   u    l   a   r   s   p    h   e   r   o    i    d   s      –     H    U    V    S    M    C   s ,    H    D    F   s    D    i   a   m   e    t   e   r   :    9    0    0   µ   m    W   a    l    l   :    3    0    0   µ   m     [    6    1     ]    N   o   z   z    l   e  -    b   a   s   e    d   :    C   o  -   a   x    i   a    l   n   o   z   z    l   e   p   r    i   n    t    i   n   g    A    l   g    i   n   a    t   e      –     B   o   v    i   n   e   c   a   r    t    i    l   a   g   e   p   r   o   g   e   n    i    t   o   r   c   e    l    l   s    D    i   a   m   e    t   e   r   :          1    0    0      –     6    0    0   µ   m     [    6    2      –     6    4     ]    A    D    S    C   s   :   a    d    i   p   o   s   e  -    d   e   r    i   v   e    d   s    t   r   o   m   a    l   c   e    l    l   s   ;    H    D    F   s   :    h   u   m   a   n    d   e   r   m   a    l    f    i    b   r   o    b    l   a   s    t   s   ;    H    U    V    E    C   s   :    h   u   m   a   n   u   m    b    i    l    i   c   a    l   v   e    i   n   e   n    d   o    t    h   e    l    i   a    l   c   e    l    l   s   ;    1    0    T    1    /    2   ;    H    U    V    E    C   s   :    h   u   m   a   n   u   m    b    i    l    i   c   a    l   v   e    i   n   e   n    d   o    t    h   e    l    i   a    l   c   e    l    l   s   ;    H    U    V    S    M    C   s   :    h   u   m   a   n   u   m    b    i    l    i   c   a    l   v   e    i   n   s   m   o   o    t    h   m   u   s   c    l   e   c   e    l    l   s   ;    H    M    V    E    C   s   :    h   u   m   a   n   m    i   c   r   o   v   a   s   c   u    l   a   r   e   n    d   o    t    h   e    l    i   a    l   c   e    l    l   s   ;    R    F    M    F   s   :   r   a    t    f   a    t   m    i   c   r   o   v   e   s   s   e    l    f   r   a   g   m   e   n    t   s   ;    H    U    V    S    M    C   s   :    h   u   m   a   n   u   m    b    i    l    i   c   a    l   v   e    i   n   s   m   o   o    t    h   m   u   s   c    l   e   c   e    l    l   s   ;    M    S    C   s   :   m   e   s   e   n   c    h   y   m   a    l   s    t   e   m   c   e    l    l   s   ;    S    M    C   s   :   s   m   o   o    t    h   m   u   s   c    l   e   c   e    l    l   s   ;    H    A   :    h   y   a    l   u   r   o   n   a   n   ;    P    E    G   :   p   o    l   y     (   e    t    h   y    l   e   n   e   g    l   y   c   o    l     )   ;    R    G    D    S   :   a   r   g    i   n    i   n   e      –    g    l   y   c    i   n   e      –    a   s   p   a   r    t    i   c   a   c    i    d      –    s   e   r    i   n   e   ;    B    i   o    L    P   :    b    i   o    l   o   g    i   c   a    l    l   a   s   e   r   p   r    i   n    t    i   n   g   ;    D    O    D   :    d   r   o   p  -   o   n  -    d   e   m   a   n    d . E. Hoch  et al . / European Journal of Cardio-Thoracic Surgery 770
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