Recent Strategies in Extrusion-based 3D Cell Printing towards Organ

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Recent Strategies in Extrusion-based 3D Cell Printing towards Organ Biofabrication Ge Gao, Byoung Soo Kim, Jinah Jang, and Dong-Woo Cho ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/ acsbiomaterials.8b00691 • Publication Date (Web): 21 Jan 2019 Downloaded from http://pubs.acs.org on January 26, 2019

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ACS Biomaterials Science & Engineering

Recent Strategies in Extrusion-based 3D Cell Printing toward Organ Biofabrication Ge Gao1, #, Byoung Soo Kim 1, #, Jinah Jang2, *, and Dong-Woo Cho 1, *

1

Department of Mechanical Engineering, Pohang University of Science and Technology, Pohang, 37673, Republic of Korea

2

Department of Creative IT Engineering, Pohang University of Science and Technology, Pohang, 37673, Republic of Korea

Prof. Dong-Woo Cho: [email protected]; Tel.: +82-054-279-5889; Prof. Jinah Jang: [email protected]; Tel.: +82-054-279-8821

Keywords: 3D cell printing; microextrusion printing; bioink; tissue engineering; organ printing.

Abstract: Reconstructing human organs is one of the ultimate goals of the medical industry. Organ printing utilizing three-dimensional (3D) cell printing technology to fabricate artificial living organ equivalents has shed light on the advancement of this field into a new era. Among three currently applied techniques (inkjet, laser-

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assisted, and extrusion-based), extrusion-based cell printing (ECP) has evoked the majority of interest due to its low cost, wide range of applicable materials, and ease of spatial and depositional controllability. Major challenges in organ reconstruction include difficulties in precisely fabricating complex structural features, creating perfusable and functional vasculatures, and mimicking biophysical and biochemical characteristics in the printed constructs. In this review, we describe the merits and limitations of ECP for organ fabrication and discuss its recent advances aimed at overcoming these challenges. In addition, we delineate the expected future techniques for printing live tissue or organ substitutes.

1. Introduction

Despite the increasing number of willing donors, the crisis of global organ shortage has grown steadily in recent years. The most recent data has indicated that 144 patients are added to the waiting list for organ transplantation every day 1. During this same 24-hour period, approximately 22 people die while waiting for a suitable donation. Reconstructing living organs from an autologous source could be a direction to solve this problem. Over the past decades, tissue engineering has


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emerged as a multidisciplinary field in the pursuit of developing artificial grafts to regenerate or replace damaged tissue and organs

2-3.

The traditional tissue

engineering approach is to seed cells onto a scaffold obtained by processing biomaterials or by decellularizing natural organs. Based on these efforts, various tissue grafts, such as skin, cartilage, bone, blood vessel, and bladder, have been engineered and validated for clinical applications 4. Although significant success has been achieved, it is difficult to apply these strategies to rebuild solid and heterogeneous organs (e.g., liver, kidney, brain, and heart) with intricate structures, multiple cell/tissue constituents, massive integrated vascular networks, and complex physiological functions.

3D cell printing is considered a promising technology due to its outstanding ability of precisely positioning multiple biomaterials and living cells in a layer-by-layer manner based on patient-specific designs acquired from medical imaging 5. To date, various types of 3D printed tissue and organs have been developed for medical applications 6-8. In addition, 3D cell printing has also been applied to establish in vitro tissue models

9

and organ-on-a-chip platforms

10-11,

providing improved

interpretations of tissue morphogenesis, pathophysiology, and pharmacological ACS Paragon Plus Environment


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efficacy or toxicity of drugs. To date, three types of 3D cell printing techniques have been widely used, including inkjet, laser-assisted, and extrusion-based techniques. Compared to the inkjet and laser-assisted techniques, extrusion-based 3D cell printing (ECP) can apply most of the fluidic biomaterials

12,

stack-up large 3D

constructs 7, and involve a high density of cells 13. Moreover, in combination with the printing of molten polymers, it is capable of building structures with controlled porosities and tunable mechanical strengths

14.

Due to these advantages, ECP is

considered the most promising technique for tissue/organ biofabrication. However, to advance the reconstruction of solid and heterogeneous organs, several current challenges should be overcome.

The first obstacle is to precisely mimic the intricate structures of native organs. An organ refers to a part of an organism that is typically self-contained and has a specific vital function. The intricacies of organs lie in not only the sophisticated shapes and architectures but also the heterogeneous distributions of different types of tissue and cells. This naturally structural and compositional definition serves to execute the biofunctions of each organ. For example, the kidney is a bean-shaped structure containing an outer renal cortex and an inner renal medulla. Spanning ACS Paragon Plus Environment

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these two divisions, approximately one million nephrons (functional units composed of renal corpuscles and convoluted renal tubules) are aligned and function to convert blood into urine. Nephrons generally have physical dimensions of 150 μm to 250 μm and contain abundant small capillaries with diameters of 5 μm to 10 μm

15.

To

reconstruct functional organs, the presence of these structures is crucial. Therefore, recapitulating the structural and mechanical features of an organ is the initial step in organ printing.

Another challenging consideration for organ printing is vascularization. The ultimate goal of tissue engineering is to create tissues that can be used as alternatives to donor tissue to replace damaged tissues and organs

16-17.

These

tissues mostly have a larger size than the diffusional limits for nutrients and oxygen 18.

The inclusion of perfusable vascular networks in engineered tissue is of

importance for ideal organ biofabrication. In addition, blood vessels in the human body play a crucial role in a variety of biological processes, including metabolism, healing, regeneration, and the immune response

19.

Thus, a strategy to generate a

functional vasculature within a large volume tissue construct is required to promote oxygen transfer, nutrient delivery, metabolic waste disposal, and immune cell ACS Paragon Plus Environment


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circulation. There is evidence that a construct (> 200 µm in thickness) without a vasculature incurs hypoxic conditions in which certain regions of the construct are deprived of an adequate oxygen supply

20-21,

which implies that living cells should

reside within a 200 μm distance from a vessel.

In addition, although 3D cell printing can position various cell types in accordance with anatomical tissue design, merely replicating the structural features is not sufficient to fabricate functional organs. Growing awareness justified that the microenvironment in bioink plays a vital role in cellular functioning and organ morphogenesis

22.

In addition to biological concerns, recapitulating the dynamic

physical microenvironments in tissue conformations could provide more biomimetic incentives to execute the unique functions of organs

23-24.

For instance, the heart

regularly beats to pump blood throughout the body, which is achieved by rhythmic contractions; the esophagus and intestines exhibit peristaltic behavior to facilitate the ingestion and digestion processes; and the brain propagates commands in the form of electrical signals through nerves to control musculoskeletal systems for body motion. Hence, it is also critical to emulate the physiology of organs in fabricated


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tissue constructs to provide cells with an environment that is analogous to their physiological environments.

In this article, we focus on the challenges that have appeared in the major steps of the fabrication process, including printing strategies and choice of biomaterials. We explain the working principles of ECP and highlight the advantages of ECP on the objectives of organ printing; then, we discuss advanced bioink design and printing approaches for recapitulating the physical shapes and structures of organs. Next, we summarize the strategies for vascularization and providing biomimetic environments via ECP techniques. Finally, we envision future directions for promoting the progress of ECP toward the ultimate goal of organ reconstruction.

2. Extrusion-based 3D cell printing (ECP) 2.1. Working principle of ECP The ECP technique combines fluid-dispensing and automated robotic systems 6. It is controlled by a computer, which precisely shows the programmed code defined by customers or computer-aided design (CAD) files converted from bioimaging data (e.g., MRI and CT scans). A 3D extrusion printer typically consists of a temperature-controlled material cartridge to which various types of trigger forces are applied to physically squeeze bioink out from a fine nozzle. Collaborative movement along the x-, y-, and z-axis enables deposition of continuous filaments on the platform of the printer. The first deposited layer provides a ACS Paragon Plus Environment


base for the next layer, and sequential layers enable the fabrication of predesigned 3D tissue structures. To achieve optimal resolution of products, researchers need to consider systematic printing parameters, including nozzle gauge, motion speed, and bioink flow rate. Notably, although using a smaller nozzle can lead to a finer printing resolution, it may negatively affect cell viability due to increased shear stress 25. Based on the driven force for extrusion, ECP can be largely divided into pneumatic and mechanical systems (piston- and screw-driven, respectively) 12, 26 (Fig. 1). Pneumatic systems employ a compressed air controller (from 5 to 800 KPa) connected with a syringe or valve configurations, which can be easily mediated at a suitable level for printing bioinks with different rheological properties (30 mPa·s to >6 x 107 mPa·s) 7, 12. Syringe systems have been widely used in this field due to their simplicity, but valve systems requiring a pulse frequency have been shown to provide greater precision 7.

Mechanical systems apply force directly to bioink using pistons or screws. Piston

configurations generally provide direct control over material flow, whereas screw configurations provide more spatial control and are beneficial for higher viscosity bioinks (> 6 × 107 mPa·s) 12 but can potentially damage encapsulated cells 27-28 because they generate large pressure drops at the nozzle. Therefore, rotating screw configurations should be carefully considered. Compared with other printing techniques, features of ECP under various conditions are summarized in Table 1.


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Figure 1. Currently available ECP systems: (A) pneumatic microextrusion including valve-free (left) and valve-based (right), (B) mechanical microextrusion including piston (left)- or screw (right)-driven microextrusion. Reproduced with permission from ref 12. Copyright 2016 Elsevier.

2.2. The necessity of ECP for organ biofabrication Although inkjet- and laser-based cell printing techniques are also widely used for tissue engineering applications, both techniques include critical limitations to meet the ultimate goal of organ biofabrication. A major drawback of the inkjet-based technique is that it is only applicable to biological materials with low viscosities (3.5-12 mPa·s) due to nozzle clogging 29-30.

For this reason, the printed material that is in a liquid form requires rapid crosslinking

prior to printing of the next layer, and during fabrication, this requires additional time, which can be time-consuming. As a result, using the inkjet-based printing technique is onerous in the fabrication of bulky 3D products with complex structural organization, notwithstanding the rapid printing speed (1-10,000 droplets per second) and high resolution (at the single cell level). Indeed, this difficulty confines their application to the fabrication of 2D or tissue-level constructs (e.g., skin and cartilage). Furthermore, the laser-assisted printing technique does not have a problem with nozzle blockage because it is a nozzle-free method 31. However, this technique is relatively cumbersome for printing complex structures with various compositions because cell-laden bioinks should be prepared in the form of individual ribbons. As a result, this technique faces formidable challenges when tissues/organs have to be printed with multiple types of cells and biomaterials 7. For this reason, researchers working on organ biofabrication may be reluctant to use the laser-based printing technique. In contrast, ECP has exhibited the greatest potential in the engineering of 3D tissues and organs. First, ECP facilitates the printing of a wide range of fluidic biomaterials (3 to 6 × 107 mPa·s), including synthetic polymers, cell aggregates, cell-laden hydrogels, and microcarriers, that provide better options for printing target organs with specific mechanical properties, cell densities,



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and microenvironments (e.g., gradients of growth factors). When combined with extruded synthetic polymers, ECP enables the fabrication of controllable porous constructs and therefore allows nutrient diffusion to support cell survival in massive constructs

32.

More

importantly, unlike the inkjet-based printing technique, ECP is capable of printing high cell densities

33-35

which is essential to promote cell self-assembly, organization, extracellular

matrix (ECM) secretion, and, eventually, tissue maturation.

However, the successful reconstruction of solid organs (e.g., heart, kidney, liver, and brain) fundamentally requires precise structural recapitulation to mimic anatomical features, the involvement of a perfusable and functional vasculature to support cell survival, and the presence of biomimetic microenvironments to encourage cellular function and tissue formation. Although ECP has been used for engineering various tissue/organ constructs 36, it is still limited in the fabrication of complex and vascularized organs with physiological functions.


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Table 1. Comparison of 3D cell printing techniques Types of 3D cell printers Inkjet

Extrusion

Laser

Refs.

Printer cost

Low

Medium

High

37

Available viscosities of materials

3.5-12 mPa·s

30 mPa·s to > 6 × 107 mPa·s

1-300 mPa·s

14, 31, 38

Gelation methods

Chemical, photocrosslinking

Chemical, photocrosslinking

39-41

Dispensing speed through a nozzle

Fast (1-10,000 droplets per second)

Slow (10 μm-50 mm/s)

Medium-fast (200-1,600 mm/s)

42-44

Preparation time

Low

Medium

High

45

Resolution

300 droplets, 50 µm wide

100 µm to millimeters wide

Microscale resolution

42, 46-48

Cell viability

>85%

40-80%

>95%

44

Cell densities

Low, 200 μm

55, 91-93

~ 150 μm

94-95

enzymatic

(> 2 months)

adhesive moieties

Proteolysis (~ 1 weeks)

cell favorable

strength

Physical and chemical

Ionic

Agarose

Physical

Chitosan

Ionic and chemical

Synthetic glycol)

Printing performance Refs.

Noncytotoxic, lacks cell

— varied

with Hydrolysis, ion exchange

molecular weight and (varied with molecular Ca2+ contents

poly(ethylene

Cytocompatibility

Physical, chemical, and Strong, ~ 25 KPa tensile Proteolysis

Tunable, Alginate

Biodegradability

Chemical

Fragile,

3-15

weight) KPa

compressive strength —

Tunable

Nondegradable Hydrolysis, proteolysis

Nondegradable

Noncytotoxic, lacks cell adhesive moieties Noncytotoxic, lacks cell adhesive moieties Noncytotoxic,

support

cell adhesion Noncytotoxic, cell adhesion

support

polymer Pluronic F127 Physical and chemical

Soft and weak

—


Cytotoxic


3. Structural recapitulation

Although 3D printing has been widely used to produce exquisite products for a variety of engineering applications, the rebuilding of sophisticated architectures with living cells has encountered more obstacles. The basic problem is that most traditional materials (e.g., metals, ceramics, and plastics) and assisting methods are no longer applicable to cell printing due to vulnerable cells and sensitive biological molecules

7, 96.

Therefore, to fabricate complex constructs that mimic the organ

structures, optimizing bioink designs and advancing extrusion strategies are necessary.

3.1. Optimized bioink designs

The main role of bioink is to escort cells and act as controllable “bricks” to build desirable complex structures. An ideal bioink for the accurate fabrication of living structures must match several desirable features, including 1) proper rheological properties (e.g., shear thinning behavior and suitable viscosity) to reduce the risk of cell damage caused by shear stress during extrusion27-28, 97; 2) high printability to dispense the structure with the designed shape; 3) tunable and cell-friendly gelation


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kinetics to provide shape fidelity with a high resolution; 4) sufficient mechanical properties to maintain the printed structures for a long period; and 5) cytocompatibility to benefit cell survival and functions. Nevertheless, such versatile materials for printing cells and constructing intricate structures have not yet been found or developed, and for this reason, researchers have improved the printing performance of bioinks while retaining their biological activity.

3.1.1. Physical blending

Hybrid bioinks composed of multiple constituents have been widely investigated to compensate for the limitations of each component because no single material is sufficiently versatile for cell printing. Physical blending has been used as the simplest method for integrating multiple biomaterials to achieve suitable printability, structural stability, and biocompatibility. For instance, although gelatin hydrogel is fugitive during incubation at 37 ℃, it can be used in combination with other biological materials that lack adequate printability to fabricate complex constructs. Rodriguez et al. developed silk-based bioinks using gelatin (1:1 silk to gelatin ratio) with glycerol as a nontoxic additive to induce physical crosslinking during printing

98.

This bioink

provided structural support for patient-specific 3D cheek geometry for soft tissue ACS Paragon Plus Environment


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reconstruction (Fig. 2a). They demonstrated that this material can maintain the shape and dimensions of a printed construct for 3 months and promote cellular infiltration and tissue integration.

Alginate has been used in various tissue engineering applications because it is biocompatible and ionically crosslinkable using calcium ions to yield robust hydrogels. However, this material does not interact with cells due to the lack of celladhesive moieties, which eventually leads to cellular apoptosis of the carried cells via anoikis 99. Hence, previous efforts have attempted to introduce other components with abundant cell-adhesive domains to ameliorate cellular affinity. Pourchet et al. formulated a mixture of 10% (w/v) bovine gelatin, 0.5% (w/v) alginate, and 2% fibrinogen to promote printability and shape fidelity

100.

The gelatin provided a

thickening property that imparted strength to the bioink after printing on a cooled substrate, while the alginate provided a scaffolding component to stabilize the printed bioink when the gelatin was solubilized at 37 °C, and the fibrinogen encouraged maturation to ensure a favorable environment for the encapsulated cells. They successfully used this bioink to fabricate complex 3D objects and humanscale ear constructs without supporting materials (Fig. 2b). ACS Paragon Plus Environment

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To improve the printability and mechanical properties of soft ECM hydrogels, synthetic materials have been used as additives due to their tunable viscosity and strength. PEG and its derivatives (e.g., PEG-diacrylate (PEGDA) and PEGmethacrylate (PEGMA)) are the most commonly used polymers in ECP

92, 101-102.

However, considering their poor ability to promote cell functions and the involvement of toxic photoinitiators 103-104, it is necessary to carefully control the concentration of each blending component at a balanced level, which ensures both printability and cytocompatibility. Other synthetic materials (e.g., Pluronic F127 and polymer poly(Nisopropylacrylamide) (pNIPAAM)) have also been used to enhance the printability of bioink 94, 105-106. Nonetheless, these thermoresponsive polymers have been reported to be toxic to cells at certain concentrations, incubation times, and molecular weights 56, 107-108.

Due to their adverse effects on cells, these compounds should be

eliminated from the final products after fabrication. For instance, Kesti et al. developed a dual-crosslinked bioink composed of methacrylated hyaluronic acid (HA-MA) and pNIPAAM

109.

The thermoresponsive nature of the HA-pNIPAAM

component provided rapid gelation and post-printing structural fidelity, thereby yielding high-resolution scaffolds. By removing pNIPAAM in a brief 4 °C washing



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step, encapsulated bovine chondrocytes maintained high cell viability (91%) within 7 days.

To ameliorate the viscosity and printability of thin, soft bioinks, nanoparticles can be added into polymeric hydrogels to provide additional fluidic friction forces

110.

Nanocellulose-based bioinks have been used to enhance physiological stability over the last decade due to their good mechanical properties and biocompatibility. Wu et al. proposed a hybrid bioink composed of alginate and cellulose nanocrystals and explored its suitability for ECP

111.

The hybrid bioink exhibited excellent shear

thinning and could be easily extruded through the nozzle (100 μm diameter) while providing good initial shape fidelity (Fig. 2c). These authors printed a liver lobulemimicking 3D honeycomb structure containing fibroblasts and hepatoma cells using the proposed bioink. The encapsulated cells showed 71% initial viability, but the viability decreased to 58.91% after 3 days, which might have resulted from the absence of cell-adhesive moieties in the bioink. Incorporating cell-favorable ECM proteins

(e.g.,

collagen,

gelatin,

and

fibrinogen)

or

peptides

(e.g.,

arginylglycylaspartic acid (RGD)) could be a potential solution. Similarly, Markstedt et al. used a mixture of nanocellulose and alginate to produce a bioink and printed ACS Paragon Plus Environment

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an anatomically shaped cartilaginous human ear

112.

Human chondrocytes in the

bioink exhibited 73 and 86% cell viability after 1 and 7 days of 3D culture, respectively.

In addition to altering the rheological properties and ability of shaping complex structures, nanoparticles can also act to provide additional biological cues (e.g., releasing growth factors and conducting electrical signals) to promote cell behavior. In a recent report, Gao et al. combined poly (ethylene glycol) dimethacrylate (PEGDMA), hydroxyapatite nanoparticles (nHAP) (200 nm) and bioactive glass (BG) (20 μm) to print bone tissues

113.

Human mesenchymal stem cells (HMSCs) were

printed with this nanocomposite bioink, resulting in uniform cell distribution and high cell viability (> 80%). The addition of nHAP significantly increased the compressive modulus of the printed structure and promoted ECM deposition and bone-related gene expression (collagen I, osteocalcin, collagen X, and MMP13).

3.1.2. Chemical modification

Chemical modification is another way to alter the characteristics of bioinks by adding functional groups to polymer chains, which enables covalent crosslinking to



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improve the stability and mechanical strength of printed constructs. One representative case is gelatin methacryloyl (GelMA), which is a gelatin derivative containing a majority of methacrylamide groups. In contrast to the reversible gelation kinetics of gelatin, GelMA undergoes photoinitiated radical polymerization, meaning that the GelMA bioink can be crosslinked under mild ultraviolet (UV) light exposure in the presence of photoinitiators. This change overcomes the thermal instability of the gelatin hydrogel while retaining its biological support for cell functions, such as cell adhesion, migration, population, and differentiation. However, it results in a low viscosity material that deprives the gelatin of printability. To overcome this problem, the use of additives can be considered. For example, Schuurman et al. added 2.4% hyaluronic acid (HA) to 20% GelMA to improve its viscosity and shape fidelity and successfully fabricated multiple-layered networks with high resolution (200 μm)

114.

The carried chondrocytes showed high viability (73 ± 2% at day 3) within the GelMA construct. In short, the addition of functional groups provides a foundation for improving the performance of bioink. Other hydrogels (e.g., HA and alginate) have been similarly modified to improve mechanical properties and control degradation kinetics 69, 115-118. However, these modifications inevitably introduce risks to cells due


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to the additional toxic chemical components and the exposure of live cells to UV light 119-120.

Therefore, the concentration of photoinitiator and the UV illumination

conditions (e.g., intensity, exposure time, and distance) should be carefully considered. Furthermore, the use of newly developed photoinitiators, such as 2′,4′,5′,7′-tetrabromofluorescein disodium salt (eosin Y), tetraacylstannanes, and lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP), which have lower toxicity compared to traditionally used agents (e.g., Irgacure 2959), can reduce the negative effects

121-124.

In addition, to avoid detrimental effects (e.g., cell damage and DNA

mutation) caused by UV irradiation (190-400 nm), another potential solution involves using materials that are curable under visible light (400-770 nm). Lim et al. suggested a system composed of ruthenium (Ru)/sodium persulfate (SPS), which is curable under exposure to visible light (wavelength: 400-450 nm), as the photoinitiator

125.

Compared to conventional methods, such as using UV light and Irgacure 2959 photoinitiator, this new system yielded significantly enhanced cell viability (67 ± 5.7% (UV) compared with 87 ± 4.8% (visible light) at day 21 after irradiation at 50 mW/cm2 for each curing treatment) when used for 3D cell printing of GelMA constructs, even at high Ru/SPS concentrations and visible light intensities.



The mechanical properties of 3D cell-printed structures can be significantly improved by chemically modifying the bioinks. One demonstrated method involved producing interpenetrating polymer networks (IPNs), referring to polymers synthesized from at least two polymer networks intertwined at the molecular level 126.

These networks cannot be separated unless chemical bonds are broken, which

yields bioink with improved mechanical properties. However, the solidification process of IPN hydrogels often involves multiple covalent crosslinking processes and is thus too slow for 3D printing. To surmount this challenge, ionic-covalent entanglement (ICE) hydrogels are currently being developed that are both physically and chemically crosslinked to rapidly form hydrogels for effective applications in 3D cell printing. For example, Bakarich et al. synthesized an ICE bioink composed of acrylamide and alginate

127

that retained a printed shape soon after printing and

allowed for a covalently crosslinked acrylamide network. The printed structure was then physically crosslinked with calcium chloride solution. The stiffness and failure stress increased from 23 to 260 kPa and 11 to 130 kPa, respectively. Similarly, Hong et al. formulated an elastomeric ICE bioink using PEGDA and alginate 91. The printed structures were able to resist mechanical stress without significant plastic


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deformation while maintaining high cell viability (75.5 ± 11.6%) over a period of 7 days (Fig. 2d). Overall, these studies showed the potential of ICE bioinks for building mechanically tough 3D live structures.

Figure 2. Optimized design of bioinks for reproducing the physical features of organs. (a) Fabrication of a 3D printed structure with a physiologically relevant geometry using a silk-gelatin bioink (upper: CT scan of patient and segmentation of cheek geometry, lower: printed cheek geometry) (reproduced with permission from ref 96; copyright 2017 Elsevier); (b) centimeter-sized complex objects using a mixture of 10% (w/v) bovine gelatin, 0.5% (w/v) alginate, and 2% (w/v) fibrinogen (left: honeycomb structure and centimeter-sized complex objective with overhanging features, middle: 3D models for printing a human ear, right: printed adult-sized ear) ACS Paragon Plus Environment


(reproduced with permission from ref 98; copyright 2017 Wiley-VCH); (c) alginateCNC bioink formulation used to fabricate a liver-mimicking construct (left: bioink formulation, right: comparison of printed structures using alginate and alginate-CNC bioink) (reproduced with permission from ref 109; copyright 2017 Elsevier); and (d) interpenetrating network (IPN) bioinks for printing; IPN bioink that was synthesized by covalently crosslinking PEG and ionically crosslinking alginate and cyclic mechanical deformation (reproduced with permission from ref 89; copyright 2015 Wiley-VCH).

3.2. Advanced printing strategies

Despite the basic challenge of printing physically biomimetic organ analogues, which is the unmet performance of bioink, exploring advanced extrusion-based cell printing methods is another feasible direction to achieve this goal. Since bioink solgel transitions are generally thermo-, PH-, photo-, and enzyme-sensitive, providing assisting external environments (e.g., physical support and an environment for accelerating gelation) during the fabrication process has been widely investigated to address the limitations in cell printing.


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3.2.1. Bath-assisted approach

Bath-assisted printing refers to dispensing bioinks into an additional reservoir, providing either physical support to avoid gravity-induced collapse or crosslinking reagents to accelerate solidification, thereby ensuring printing quality and enabling stack-up fabrication

128-129.

To physically maintain a 3D cell-printed construct, the

bath should be biocompatible to benefit cell viability and have appropriate rheological features to allow for bioink extrusion and writing. In addition, the simple removal of bath materials is very important for obtaining the desired depositions. Hinton et al. used a gelatin slurry as the bath material by shattering 4.5% (w/v) gelled gelatin with pulse geared mechanical blending (Fig. 3a) 130. The resultant gelatin slurry exhibited Bingham plastic properties during printing, resisting the yielding until a threshold shear force is reached. Due to the thermally reversible features of gelatin, the bath material could be easily removed by incubation at 37 °C, leaving the intact 3D printed construct.

Carbomer (trademarked as Carbopol) hydrogel is another candidate for bathassisted printing due to its ability to remain solid at low concentrations and provide recoverable yielding in the granular state. Tapomoy et al. evaluated Carbopol ETD ACS Paragon Plus Environment


2020 performance at various concentrations ranging from 0.05-1% (w/v) for 3D printing of various materials, including silicones, hydrogels, colloids, and living cells 131.

The bath can also provide a reservoir for crosslinking agents. Crosslinking agents typically associate with bioink through specific gelation mechanisms (e.g., ionic or enzymatic), allowing rapid transformation from fluidic bioink to stable solid filaments. This concept has been used to print droplets using jetting-based techniques (e.g., inkjet or laser-assisted cell printing) 132-134, but with liquid crosslink inducers, the extruded bioink often failed to precipitate if the density was less than the reservoir solution. Therefore, mixing crosslinking reagents within the gel reservoir provides an advanced method to avoid buoyancy effects. For example, 11 mM CaCl2 and 0.1 U/ml thrombin have been included in a gelatin microparticle bath to benefit the printing of 4% (w/v) alginate and 10 mg/ml fibrinogen bioink, respectively, achieving resolutions up to 200 μm

130.

Due to ionic and enzymatic

crosslinking mechanisms, alginate and fibrinogen bioinks can transit to the gel phase in an environment with a selected concentration of crosslinker. This technique allows for the successful fabrication of complex 3D structures (e.g., spiral, branched, ACS Paragon Plus Environment

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perfused, and brain-/heart-mimicking shapes) using alginate bioink, which has been shown to produce stiff gel structures in the bath with a high concentration of calcium. Although the bath-assisted approach helps print 3D constructs with intricate internal and external architectures using cell-laden bioinks, the preparation of suitable bath materials adds significant burdens to already laborious fabrication and sterilization considerations. In addition, studies have demonstrated that the concentration of calcium has a profound influence on cellular survival and proliferation

135.

Although

the use of bioink in the 3D cell printing process can protect against the adverse effect of high concentrations of calcium on cellular activities 136, it is still important to control the content of ambient calcium. Several reports have demonstrated well-maintained cell viability and populations when 100 mM of calcium solution was used for crosslinking alginate-based bioink

137.

In addition to concentration, the prolonged

time of cell exposure to the calcium environment is also detrimental to cells

138.

Therefore, it is crucial to control the concentration of detrimental crosslinking reagents and accelerate the printing process to retain cell activities.

3.2.2. Aerosol spraying



Crosslinking cues can be supplied in the form of aerosols to induce gelation of printed bioinks. Compared with an aqueous bath, spraying aerosolized crosslinking reagents prevents suspension of printed struts within the liquid medium due to buoyancy. However, the bioink considered for aerosol treatment should quickly crosslink so that it can be printed as firm gel filaments to ensure printing resolution and enable complex structure fabrication 139. Thus, alginate is a good candidate due to its almost instantaneous gelation when treated with calcium ions. Ahn et al. fabricated a 3D cell-laden scaffold utilizing 3.5% (w/w) alginate and a crosslinking aerosol produced by fuming 2% (w/w) calcium chloride solution with an ultrasonic humidifier (Fig. 3b) 140. They used a 310 μm diameter printing needle to print a thick construct (4.5 mm height) with a homogeneous pore size (435 ± 32 μm) and fine struts (355 ± 28 μm). Notably, the encapsulated preosteoblasts (MC3T3-E1) retained high cell viability (85%) after printing, suggesting that the chosen fabrication method was not significantly detrimental to cells.

3.2.3. In situ photopolymerization

The photocurability of hydrogels containing photoinitiators has inspired researchers to integrate UV illumination with ECP techniques. Polymerization is ACS Paragon Plus Environment

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triggered by exposing bioinks, including photocrosslinkable precursors, to UV light, which results in material stiffening. The stabilization of the dispensed threads can be enhanced by adjusting the UV intensity and irradiation time, as well as the bioink formulation. Ideally, bioinks should crosslink and form stable gels immediately prior to deposition, ensuring printability. Ouyang et al. demonstrated in situ crosslinking using a photopermeable capillary as a reaction station to apply UV light to the bioink, inducing crosslinking during deposition (Fig. 3c)

141.

Various photocurable bioinks

have been analyzed to investigate the versatility of this printing technique, including methacrylated hyaluronic acid (HA-MA), gelatin methacryloyl (GelMA), polyethylene glycol diacrylate (PEGDA), and norbornene-functionalized hyaluronic acid. This method also showed high cell viability (95% after printing) and achieved very high resolution (< 60 μm) depending on the capillary size. However, UV treatment must be controlled at cell-safe levels to minimize negative UV exposure

142-145,

and the

resulting slow printing speed significantly extended the overall printing time for stabilizing bioink fibers.

3.2.4. Thermally controlled printing



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A large group of biomaterials relies on thermosensitive sol-gel transitions (e.g., collagen, gelatin, Matrigel, agarose, and Pluronic F127 (PF127)). Combining heating or cooling apparatuses into print heads or deposition platforms exposes bioinks to different environmental temperatures within the printing process, fulfilling various printing purposes, such as construct fabrication and temporary fabrication support. For example, collagen and Matrigel can be irreversibly crosslinked at 24-37 °C 147.

gel

146-

However, this thermal gelation typically requires several to dozens of minutes to form a stable 148,

which is too slow to provide sufficient shape fidelity for stacking-up complex

3D constructs. To overcome this drawback, tuning the printing temperature within the sol-gel transition range can be an effective solution. Ahn et al. developed a 3D cell printing system incorporating thermal modules to heat both the printing cartridge and dispensing plate and constructed a precisely stacked cell-laden bioink (Fig. 3d) 149.

The bioink was composed of decellularized ECM (dECM) from the skin

containing collagen components, which exhibited enhanced elastic modulus with increased temperature. Compared with normal printing conditions (23 °C), the warm environment (36 °C) provided significantly improved printability and higher accuracy of the fabricated 7.5-mm-high hollow square structure (25 layers) as well as a liver-


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shaped 3D construct using 2% dECM bioink. In contrast, cooling systems have been employed to 3D print bioinks that gel under low temperatures, such as agarose (2630 °C) and gelatin ( 90% for each cell type) 7 days after printing. ACS Paragon Plus Environment


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An even more important merit than the mechanical reinforcement of multipledispensed combinations is the ability to mimic intricate architectures and heterogeneous compositions of biological tissues and organs. Kang et al. configured an integrated tissue-organ printer (ITOP) containing multiple dispensers and fabricated human scale tissue constructs using PCL, PF127, and bioinks encapsulating different tissue-specific cells (Fig. 3e) 151. Similar to Shim et al. 14, PCL and PF127 were used as body and sacrificial materials to build the porous framework with complex shapes, and the bioink composed of gelatin (35-45 mg/ml), fibrinogen (20-30 mg/ml), HA (3 mg/ml), and glycerol (10% v/v) was applied to deliver relevant cells. They successfully produced tissue analogues, including bone, cartilage, and skeletal muscle. Reconstruction of other tissue types, including ear

60,

adipose

32,

and skin 41, have been reported elsewhere using the same approach.

In contrast to parallel organization, multiple needles can be positioned coaxially so that the spatial distribution of different materials converges into a single thread to enhance printability and produce fibers with complex structures or heterogeneous compositions. Colosi et al. used a dual concentric nozzle to produce a blended GelMA (4.5% w/v) and alginate (4% w/v) bioink in the core and 0.3 M CaCl2 solution ACS Paragon Plus Environment

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in the shell

152

(Fig. 3f). When the core and shell materials came into contact at the

outlet of the coupled nozzles, the alginate in the bioink was ionically gelled by exposure to Ca2+ ions, forming robust 150 μm diameter cylinders and providing excellent printability for stacking a thick structure (1 mm height). The GelMA was subsequently crosslinked by UV treatment, and the alginate gradually dissolved within culture media over 10 days. Because of the good cytocompatibility of GelMA, the encapsulated human umbilical vein endothelial cells (HUVECs) showed improved cell migration and organization.

Parallel coaxial nozzles can be used to produce multiple unidirectionally aligned fibers. Costantini et al. customized a microfluidic printing head connected to a coaxial extruder to fabricate multicellular constructs containing two different cell types (muscle progenitor cells and fibroblasts) that were printed using a poly (ethylene glycol) (PEG)-fibrinogen-alginate mixture

153

. Conspicuous compartmentalization of

encapsulated cells and the formation of parallel-aligned long-range myotubes demonstrated the potential of this printing strategy for skeletal muscle tissue engineering.



Figure 3. Representative examples of evolved ECP strategies. (a) A schematic of the printing of a complex 3D structure in a supporting bath made from a gelatin slurry followed by bath removal, which results in the fabrication of a chicken heart model (from left to right on the bottom: 3D CAD model, fluorescence image, and dark-field image; scale bar: 1 cm) (reproduced with permission from ref 128.; copyright 2015 American Assoc. Advancement Science); (b) a schematic showing the advantage of

in situ crosslinking for the printing of delicate structures compared with pre-/postcrosslinking methods (reproduced with permission from ref 139; copyright 2017 Wiley-VCH); (c) a schematic of alginate-based bioink printing supplemented with an aerosol-spraying process, producing semi-crosslinked struts (orange: nonACS Paragon Plus Environment

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crosslinked, purple: crosslinked by exposure to CaCl2 aerosol) (reproduced with permission from ref 138; copyright 2012 American Chemical Society); (d) a conceptual diagram of thermally controlled 3D cell printing applying bottom and top heating and the resultant fabrication of a liver-shaped construct (left: top view, right: side view) (reproduced with permission from ref 147; copyright 2017 Nature Publishing Group); (e) a human-scale ear tissue construct fabricated through the use of multiple nozzles equipped with an ITOP system (reproduced with permission from ref 149; copyright 2016 Nature Publishing Group); and (f) a schematic of coaxial nozzle-enabled 3D cell printing, where the CaCl2 (blue dots) in the shell are simultaneously printed with the bioink (GelMA (red), alginate (green), photoinitiator, and cells (gray)) in the core, which induces gelation of the alginate component (reproduced with permission from ref 150; copyright 2016 Wiley-VCH).

4. Vascularization strategies in ECP

An optimal vascular network in an engineered tissue needs to possess several characteristics. One of the key tasks of a vascular network is to provide all of the cells in a tissue with sufficient nutrients and oxygen. For this purpose, a vascular network organized as a complex vascular tree is required. In addition, the vascular ACS Paragon Plus Environment


network should act as a barrier between the vessel lumen and surrounding tissue while controlling the passage of materials and the transit of white blood cells into and out of the bloodstream and endothelium Finally, the construct should be easily connected to the vasculature of the patient through microsurgery, which requires vascular structures with a diameter of several hundred micrometers. Currently, various printing strategies for the vascularization of engineered tissues have been explored based on synthetic polymers, biomaterials, cells, and angiogenic growth factors. In this section, we cover these types of printing strategies based on two key aspects, indirect and direct printing techniques.

4.1. Extrusion of fugitive molds

In the indirect extrusion approach, many research efforts on prevascularized tissue engineering have relied on the spontaneous organization of endothelialrelated cells to form vascular networks on scaffolds 154-155. These studies have been successful in demonstrating the potential of adding a vascular network to engineered tissues, but only relying on uncontrolled angiogenesis results in the random organization of the vascular network that does not provide clear locations for surgical anastomosis, leading to a delay in network perfusion. Furthermore, depending only ACS Paragon Plus Environment

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on natural organization on a large scale is limited due to the relatively slow ingrowth process, which is unfavorable to cell activity during early stage postimplantation 156. Therefore, for realistic organ-scale biofabrication in the future, an in vitro perfusable hierarchical vascular network must be considered in compliance with vascular anatomical features, which means that more sophisticated and advanced printing approaches featuring greater flexibility, speed, and versatility are required.

To fabricate perfusable vascular structures based on the indirect method, bulk hydrogels, such as alginate, agarose, gelatin, or Pluronic F127, can be removed after printing by thermal or chemical treatment, leaving behind the vascular pattern once the sacrificial hydrogel is removed 26. Endothelial cells should be gently seeded through the channel. For example, Wu et al. developed 200-600 μm diameter perfusable hierarchical microvascular networks based on omnidirectional fabrication by printing 23% (w/w) Pluronic F127 (PF127) within a reservoir containing 23% (w/w) PF127 diacrylate (PF127-DA) gel (Fig. 4a)

157.

The structures were irradiated with

UV light (365 nm) for 5 min after fabrication to photocrosslink PF127-DA, and the fugitive PF127 was removed to create perfusable channels. Using a similar methodology, Homan et al. also fabricated a perfusable chip with 3D convoluted ACS Paragon Plus Environment


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renal proximal tubules by casting a cell-laden gelatin-fibrinogen hydrogel into a convoluted cylinder printed network using PF127 (Fig. 4b), followed by evacuation of the liquefied PF127 and seeding of proximal tubule epithelial cells 158. In addition to PF127 as a sacrificial hydrogel, gelatin and agarose have also been used as effective templates for the formation of perfusable channels (Fig. 4c)

156.

Based on

this methodology, Massa et al. fabricated perfusable vascularized structures using a sacrificial agarose fiber to study drug toxicity

159.

This microchannel approach

incorporated an endothelial layer using the identical methodology of previous approaches. The construct containing the endothelial layer showed increased cell viability, which demonstrated the barrier role of the induced endothelial cell layer. However, because this indirect extrusion requires additional procedures, such as cell seeding, it sometimes results in low efficacy of cell seeding. In addition, chemical processes to remove the sacrificial hydrogel might be harmful to cell function

54.

Therefore, further technological improvements and proofs are necessary.

4.2. Direct printing of microfluidic channels

For the direct approach, tubular structures containing an endothelial layer are directly extruded onto the desired vascular architecture by means of a coaxial cell ACS Paragon Plus Environment

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printing technique. After printing, the cell-laden bioink is solidified so that the cells in the bioink are exposed to a favorable environment that eventually leads to cell proliferation and tissue maturation. For the first time, Zhang et al. directly bioprinted vessel-like cellular microfluidic channels in the form of hollow filaments, where the hydrogel was extruded through a feed tube into the space between the outer and inner tubes

80.

After contacting the crosslinking solution, gelation occurred

instantaneously and resulted in tubular constructs (Fig. 4d). Their group further demonstrated vascular tissue fabrication by adding human umbilical vein smooth muscle cells, and perfusion and permeability experiments were successfully performed using these constructs. Cell viability was maintained at 84% after 7 days of perfusion, thus supporting their potential as functioning blood vessels. Based on the identical methodology, Jia et al. used a coaxial nozzle to produce blended GelMA (7% w/v), alginate (3% w/v), and PEGTA (2% w/v) bioink in the core (Fig. 4e)

160.

When these materials encountered each other at the outlet of coupled nozzles, the alginate in the bioink was ionically gelled by exposure to Ca2+ ions, forming perfusable conduits with an outer diameter of 800 μm and a wall thickness of 110 μm and providing excellent printability for stacking a thick structure (10 layers). The



GelMA was subsequently crosslinked by UV treatment, and the alginate gradually dissolved within culture media over 10 days. Because of the good cytocompatibility of GelMA, the encapsulated human umbilical vein endothelial cells showed improved cell migration and organization. More recently, Gao et al. engineered an endothelial progenitor cell (EPC)/atorvastatin-loaded poly(lactic-co-glycolic) acid microsphereladen bio-blood vessel using a 3D coaxial cell printing technique for ischemic disease therapy

161.

In this work, vascular-derived ECM (VdECM) bioink was

developed to provide a favorable microenvironment. Furthermore, this bioink was mixed with alginate for 3D coaxial cell printing; the alginate in the bioink played a key role in the initial crosslinking prior to the thermal crosslinking of VdECM, which led to the stabilization of the tubular structure during printing. In vivo studies showed improved survival and differentiation of EPCs, enhanced neovascularization rate, and salvaging of the ischemic limbs.

Direct extrusion has been the main focus for engineering vessel structures with complex structural integrity. However, several concerns must be overcome in terms of technological and material issues. For example, bioprinting complex patterns of vascular networks at multiple scale ranges is still far from being achieved, and the ACS Paragon Plus Environment

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aforementioned printing technologies fall short of producing an intact capillary network at the single cell level due to their limited resolution 156. Further studies might focus on the improvement of nozzles (termed a print head), such as studies by Mueller et al.; they developed multiple microscale nozzles within a printing head and showed its capability to print fine patterns with bioinks 162. In addition, tissue-specific bioink materials need to be formulated with vascularization-supportive bioactive materials and be used under high fidelity to allow rapid fabrication of vascularized thick tissues at clinically relevant volumes. Overall, further technological and material advances should be considered in the future.



Figure 4. Representative examples that show indirect and direct extrusion techniques for vascularization. (a) The concept of omnidirectional printing of 3D microvascular networks within a hydrogel reservoir (reproduced with permission from ref 155; copyright 2011 Wiley-VCH); (b) 3D convoluted renal proximal tubule on chip; a schematic of the indirect fabrication process using PF127 (left); a 3D rendering of the printed convoluted proximal tubule acquired by confocal microscope (right) (reproduced with permission from ref 156; copyright 2016 Nature Publishing ACS Paragon Plus Environment

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Group); (c) a schematic that describes indirect extrusion using agarose (reproduced with permission from ref 154; copyright 2017 Elsevier); (d) coaxial 3D cell printing; (i) a schematic of vascular-like structures of the coaxial extrusion process; (ii) an image of bioprinted alginate microfluidic channels; (iii) cell viability of alginate vascular conduits (reproduced with permission from ref 78; copyright 2013 The American Society of Mechanical Engineers); and (e) coaxial bioprinting of a 3D construct (10-layer stacking) with perfusable hollow tubes using a blended bioink composed of alginate, GelMA, and PEGTA, which allow the encapsulated vascular cells to achieve early maturation of vessel tissue on day 21 (green: α-SMA, red: CD31, blue: DAPI) (reproduced with permission from ref 158; copyright 2016 Elsevier).

5. Biomimetic microenvironments

Many efforts have attempted to mimic some aspects of native ECM using naturally derived purified biomaterials

163.

Some growth factors and cytokines have

been supplemented as additives in bioinks to emulate certain advantageous features of the native ECM 154-155. Despite these efforts, human organs are more structurally complex than expected because they are composed of many undiscovered ACS Paragon Plus Environment


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supportive factors and functional bioactive components secreted by resident cells. Moreover, ECM proteins and signaling molecules in native organs vary for each tissue and organ/tissue specificity; thus, it is almost impossible to emulate the intricate microenvironment in the native organs by current engineering approaches. This suggests that tissue-specific bioink materials might need to be ideally formulated with tissue-specific growth-directing bioactive materials for potential 3D functional organ biofabrication. More recently, the concept of mimicking the dynamic motions of natural counterparts has been further suggested to emulate the authentic physiology of target tissues and organs beyond biochemical mimicry. In this section, we

describe

recent

research

trends

in

providing

cells

with

improved

microenvironments in terms of tissue specificity and dynamic motion.

5.1. The tissue-specific microenvironment

Utilizing the ECM itself as a source of bioink has been popular in the field 164-165. Several studies have regarded decellularized ECM (dECM) as a potential bioink source that is able to provide a tissue-specific microenvironment

166-167.

As a

pioneering study, Pati and coworkers initiated the concept of tissue-specific dECM bioink and suggested its potential for constructing analogous tissue via 3D cell ACS Paragon Plus Environment

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printing techniques (Fig. 5a)

167.

To avoid adverse immunogenicity, adipose-,

cartilage-, and heart-relevant dECM was accordingly derived from porcine tissues following different decellularization protocols to remove cellular and nuclear components. Biochemical assays indicated that DNA content was sufficiently reduced (< 50 ng/mg ECM dry weight), while key protein components, such as collagen and glycosaminoglycans, were largely preserved. After acidic pepsin digestion, the resultant dECM bioinks exhibited desirable rheological properties, including shear-thinning behavior, thermal-gelation kinetics, and sufficient shaperetention modulus, after crosslinking. They successfully fabricated 3D constructs using a multiple nozzle printing strategy combining polycaprolactone (PCL) and cellladen dECM bioinks. Compared with type I collagen, cells encapsulated in three types of dECM bioinks not only exhibited higher cell viability but also better tissuerelevant gene expression, reflecting superior biofunctionality in supporting cell survival as well as directing cell differentiation into tissue-specific lineages (Fig. 5b). Based on these observations, they developed a 3D cell-printed dome-shaped structure composed of a dECM bioink and PCL for breast reconstruction

32.

The

human adipose-derived stem cells within the dECM were printed onto the internal



pores of the supportive PCL construct (Fig. 5c-i). After 2 weeks in culture, the adipogenic markers PPARγ and COL IV were identified without the aid of any supplemented adipogenic factors (Fig. 5-ii). This observation indicates that the tissue-specific ECM in dECM can influence stem cell lineage and differentiation. For

in vivo studies, the 3D cell-printed dome-shaped construct was implanted at a subcutaneous site of nude mice and was observed for 12 weeks. This result showed that host cell infiltration into the construct and construct-enabled tissue formation were improved compared to nonprinted decellularized adipose tissue gel (Fig. 5c-iii, iv). More recently, Jang et al. demonstrated a 3D-printed prevascularized tissue construct using micropatterning of dual stem cell-laden decellularized heart matrix bioink. It ameliorated cardiac function and cellular infiltration into infarct areas and reduced cardiac hypertrophy and fibrosis (Fig. 5d) 155, 168. Following these outcomes, several more bioinks have been formulated, demonstrating the potential of tissuespecific dECM bioinks in regulating cellular functions and thus regeneration of various tissues, including muscle 169, liver 170, bone 171, and vascular tissues 161.

However, similar to other natural hydrogels, the main disadvantage of dECM bioink is weak printability, which suggests combinations with advanced printing ACS Paragon Plus Environment

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techniques (introduced earlier) for functional 3D organ biofabrication. In addition, due to the variety of tissue in their compositional, mechanical, and biochemical properties, the decellularization protocols vary depending on the target, indicating that additional efforts are also required to extend the library of dECM bioinks.

Figure 5. Representative examples that show tissue-specific bioinks and their applications. (a) Decellularization of cartilage, heart, and adipose tissues; (i) images of decellularized tissues and histological analysis compared with native tissues; (ii) ACS Paragon Plus Environment


biochemical analysis of developed bioinks; (iii) fabricated cell-laden constructs with dECM bioinks; (b) evaluation of differentiation into tissue-specific lineages and structural maturation of cells; immunofluorescence images showing chondrogenic differentiation of hTMSCs in (i) COL and (ii) cartilage dECM constructs showing COL type II (red), cell nuclei (DAPI, blue), and F-actin (green) (scale bar: 50 um); structural maturation of myoblasts in (iii) COL and (iv) heart dECM constructs showing Myh7 (red) and cell nuclei (DAPI, blue) (scale bar, 200 um); adipogenic differentiation of hASCs in (v) COL and (vi) adipose dECM showing PPARγ (red), COL IV (green), and cell nuclei (DAPI, blue) (scale bar, 50 um) (reproduced with permission from ref 165; copyright 2014 Nature Publishing Group); (c) fabrication of dome-shaped construct (i); evaluation of adipogenic differentiation without exogenous factors with adipogenic markers (PPARγ (red), COL IV (green), and cell nuclei (blue)) in vitro (ii); H&E and immunohistochemical images (cell nuclei, blue; collagen IV, green; PPARγ, red) of the implanted constructs at 12 weeks postimplantation (iii), (iv) in turn (reproduced with permission from ref 30; copyright 2015 Elsevier); (d) a schematic of a prevascularized stem cell patch including multiple cell-laden bioinks and a supporting PCL polymer (i); images of the implanted


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patch in the epicardium for each group (ii) (reproduced with permission from ref 153; copyright 2017 Elsevier).

5.2. The dynamic microenvironment

The concept of 4D bioprinting has been proposed to simulate natural dynamics. 4D bioprinting is based on 3D printing but further integrates programmable shape transformation induced by external stimuli (e.g., temperature, humidity, magnetism, electricity, light, etc.) or cellular activity over time (e.g., cell contraction, cell migration, etc.)

23, 172.

A significant amount of stimulus-responsive materials, which are also

called smart materials, have been developed for smart textiles, autonomous robotics, biomedical devices, drug delivery, and tissue engineering

173-176.

However, few of

these materials have been used as ECP bioinks due to additional requirements for successful printing, such as appropriate rheological parameters, suitable physical stability, and strong biocompatibility. For example, Jamal et al. reported a self-folding hydrogel scaffold composed of two PEG layers with different molecular weights (4000 and 10000 MW)

177.

The shape transformation was driven by the distinct

swelling ratio of the two PEG layers in aqueous surroundings, thereby curving the scaffolds (Fig. 6a). The cell scaffold structurally transitioned into cylinders of different ACS Paragon Plus Environment


radii in the presence of water, and the process was not detrimental to the encapsulated insulin- secreting β-TC-5 cells, which was indicated by high cell viability and insulin production over eight weeks. However, current fabrication of these scaffolds involves photolithography rather than extrusion because of the weak printability of the applied PEG.

Furthermore, Bakarich suggested an ICE hydrogel composed of 4% (w/v) alginate and 10-20% (w/v) pNIPAAM, which is mechanically robust, thermally actuating, and particularly suitable for 3D extrusion

127.

By changing the amount of

pNIPAAM (thermally responsive to volume change), the hydrogel exhibited reversible length alteration of 41-49% as the temperature shifted between 20 °C and 60 °C. A smart valve incorporating this feature was successfully printed to control water flow by automatically closing upon exposure to hot water and opening in cold water. Similarly, Gladman et al. designed a printable hydrogel composed of synthetic hectorite clay, nanofibrillated cellulose, and N,N-dimethylacrylamide or Nisopropylacrylamide (NIPAM) monomers, thereby achieving shape-morphing systems under environmental stimuli (Fig. 6b) and incorporating anisotropic swelling properties by orientating cellulose fibrils

24.

During the printing process, fibrils could


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be forced to align along the longitudinal direction under strong shear stress controlled by printing parameters (e.g., nozzle size, flow rate, and printing speed). By manipulating this property, they programmatically fabricated plant-inspired architectures, such as flowers, bracts, and leaves, which changed shape upon immersion in water. In particular, when the petals were composed of bilayers with 90°/0° or 45°/45° angle arrangements, the printed structure could precisely mimic the opening and closing action of flowers. However, although these bioinks and combined printing technologies have opened new avenues to emulate natural dynamics, the inclusion of living cells in the printing process has not been reported. Thus, significant work remains before achieving authentic 4D cell printing.

Another potential direction is the utilization of cell traction force (CTF) to dictate 3D structures for predefined shapes. Shigetomi et al. proposed the concept of cell origami to harness living cells as a self-folding driving force to create a diverse range of 3D cell-laden microstructures 178. They verified the concept with a model in which cells were seeded across two microplates attached to a piece of glass substrate where microplate folding could be triggered by CTF after detaching the plates from the substrate. Following this principle, assemblies of 2D microplates could be used ACS Paragon Plus Environment


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to create various 3D cell-laden structures by appropriately designed microplate geometric patterns, such as hollow cubes, dodecahedrons, and tubes. Although this methodology is distinct from the concept of 3D cell printing, it reminds us that the cell itself could be utilized to emulate natural dynamics. Traditional 3D cell printing is merely applied as an engineering tool to shape the desired structures, yet tissue generation and maturation are anticipated upon intrinsic cell functionalities such as cell migration, alignment, and aggregation. However, it is important to note that the printing process is also capable of defining intrinsic features of a 3D construct (e.g., cell alignment, fiber orientation, multiple bioink orchestration, etc.), which might provide additional signals facilitating tissue development. For example, Mozetic et al. fabricated a muscular tissue precursor by controlling C2C12 cell alignment relying on only the printing procedure

179.

In another case, McClendon et al. directed the

circumferential orientation of smooth muscle cells by trapping them in a tubular hydrogel structure with peptide amphiphilic nanofibers patterned accordingly, thereby mimicking the anatomical features of blood vessels

180.

Therefore, except

for the engineering role, it is also possible and important to use 3D printing techniques for exploring the inherent potential of cells. In the future, beyond these


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structural

and

physical

features

of

engineered

tissues/organs,

desirable

physiological functions should be taken into consideration. Examples of some tissues/organs targeted by ECP techniques are summarized in Table 3.


ACS Biomaterials Science & Engineering 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46

Page 54 of 81

Table. 3 Examples of 3D extrusion-based cell-printed tissues/organs Tissues/organs Skin

Cell source Keratinocytes and

Bioink material Gelatin, alginate, and fibrinogen

Methods Extrusion

fibroblasts Keratinocytes and

Collagen type I Porcine skin-derived extracellular matrix

fibroblasts Vessels

Stratified epidermis and ECM secretion in

Refs. 100

dermis

fibroblasts Keratinocytes and

Achievements in physiological function

Extrusion and


inkjet

dermis

Extrusion and


inkjet

dermis, and improved barrier function

Endothelial progenitor

Porcine aorta-derived extracellular

Coaxial

Formation of endothelium and improved

cells

matrix

extrusion

neovascularization in vivo

Endothelial cells

PEG-DA, Matrigel, fibrin gel

Extrusion

Formation of endothelium and sustained

41

181

161

162

metabolic function Brain

Endothelial cells

Collagen type I

Extrusion

microvasculature Cardiac

Formation of engineered brain

9

microvasculature with barrier function Cardiac progenitor cells

Porcine heart-derived extracellular

and mesenchymal stem

matrix

Extrusion

Promoted vascularization and

155

neovascularization and decreased fibrosis of

cells

injured myocardium in vivo

Osteochondral

Mesenchymal stem cells

Atelocollagen and hyaluronic acid

Extrusion

Enhanced bone formation in vivo

182

Skeletal muscle

Myoblasts

Porcine muscle-derived extracellular

Extrusion

Myotube formation and contraction in

169

matrix Ear

response to electrical stimulation in vitro

Chondrocytes

Gelatin, fibrinogen, hyaluronic acid

Extrusion

Formation of cartilage tissue in vivo

151

Chondrocytes and

Alginate

Extrusion

Gene expression for chondrogenesis and

60

adipose-derived stem

adipogenesis in vitro

cells Adipose

Adipose-derived stem

Human adipose-derived extracellular

cells

matrix

Extrusion

Vascularized adipose tissue regeneration in vivo


32

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Figure 6. Representative examples of advanced bioink designs for providing dynamic environments to the printed constructs. (a) (top) A schematic of the fabrication of a bilayer structure composed of two types of PEG with different swelling ratios using photocrosslinking steps, (bottom left) a deconvoluted fluorescent image of a self-folded bilayer with fibroblasts stained with Hoechst (blue) and calcein AM (green) in the inner and outer layer, respectively, and (bottom left) the insulin production of β-TC-6 cells cultured within a self-folded micropatterned hydrogel (blue



bar: basal insulin secretion (5.5 mM glucose), red bar: stimulated insulin secretion (16.5 mM glucose)) (reproduced with permission from ref 174; copyright 2012 Elsevier); (b) (top) a schematic of shear stress-induced orientation of cellulose fibers during the 3D extrusion process and the resultant effect on anisotropic stiffness (E) and swelling strain (α), and (bottom) a flower-shaped construct composed of 90°/0° (left) and -45°/45° (right) bilayers oriented with respect to the long axis of each petal, with time-lapse sequences of the flowers during the swelling process (bottom panel) (scale bars, 5 mm, inset = 2.5 mm) (reproduced with permission from ref 23; copyright 2016 Nature Publishing Group).

6. Conclusion and future perspectives

3D cell printing is a promising tissue engineering method for reconstructing human organs. In particular, the ECP technique has attracted great attention due to its distinct abilities in applying a wide range of biological materials, scaling-up 3D structures, and carrying a high number of cells, which are basic requirements for recreating tissues and organs. However, to advance the reconstruction of complex, vascularized organs, several prerequisites are needed, including the precise reproduction of organ structures, the involvement of perfusable vasculatures, and ACS Paragon Plus Environment

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the presence of biomimetic cellular habitats. Recent studies have attempted to overcome these obstacles by optimizing bioink designs and have suggested advanced printing strategies. This review provides a comprehensive overview of the recent efforts in regard to this progress, as well as their potentials, achievements, limitations, and future directions.

However, despite these remarkable advances, many challenges remain to be solved. First, the fabrication time should be minimized for printing living organ substitutes with clinically relevant sizes for the purpose of transplantation. Prolonged time can lead to an increased risk of nozzle clotting and cell death

183.

It has been

reported that a bioprinter needs to operate for several hours to fabricate scalable organs, such as a mouse liver with 1.3 × 108 cell/g

184.

Considering the potential

damage to cells during long-term fabrication, one strategy to accelerate the fabrication is the printing of cell aggregates and strands. Based on the ability of selfassembly, it may be possible to obtain intact organs through tissue fusion. Some researchers have successfully fabricated human-sized tissues and organs, thus demonstrating the feasibility of this idea 185. However, considering the large size and high cell density in tissue units, minimizing the hypoxic conditions in the central ACS Paragon Plus Environment


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region to support cell survival is a challenging issue. In addition, more manufacturing and biological information is necessary to produce tissue units with appropriate shapes and dimensions adaptable for use in printing.

Improving the printing resolution is another future task for organ biofabrication. Although printing complex-shaped organs (e.g., heart, brain, and ear) has been achieved using advanced ECP, the current resolution (> 100 μm) is still insufficient to replicate the structural features of organs at the anatomical level. To compensate for the low resolution of ECP, direct printing of small premature tissues, which are engineered in vitro to present physiological structures of functional organ units, can be regarded as a strategic approach. Morizane et al. reported the development of functional nephron organoids derived from human pluripotent stem cells (dimensions of 100 μm to 200 μm)

186-187.

Considering the

ability of ECP to print cell aggregates, it could be possible to use these mature organoids to fabricate complex 3D constructs. Therefore, accurately positioning the organoids without damaging their function and structural integrity can be a future direction of ECP-based organ biofabrication.

In addition, advanced bioimaging techniques and processing software are also indispensable for the future of organ printing. Indeed, MRI and CT scanning have greatly helped to visualize the outline and internal structure of patient tissues and organs. Current software is able to convert these data to numerical control (NC)


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languages that could be used to command a 3D printer to replicate the template. However, human organs are not merely structural constituents but live collections of different types of tissues. To reconstruct organ equivalents, the indispensable information is not only the structural profiles but also the precise positions of the involved cells and tissues. Although some advanced bioimaging approaches could help to reflect the maps of certain tissues (e.g., angiography could specifically visualize blood vessels) and anatomical information could provide the general blueprint of each organ, the final objective of organ printing is to rebuild patientspecific organs with unique features.

Moreover, developing a more sophisticated 3D cell printer is of the utmost importance in propelling the advancement of organ printing. Most currently used printing devices are merely equipped with a single element (i.e., an inkjet, extrusion, or laser-based module). As mentioned at the beginning of this review, each of these techniques has its own limitations. Therefore, a cell printer integrating two or more techniques could shorten the path toward organ fabrication. In a recent example, Kim et al. demonstrated the feasibility of integrating jetting (inkjet) and extrusion modules into an integrated composite tissue/organ building system. Based on this ACS Paragon Plus Environment


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device, a biomimetic full human skin model composed of functional dermis and epidermis was successfully fabricated 41.

Furthermore, except for engineering tissue and organ constructs at the in vitro stage, in situ printing, where living tissues and organs are directly printed into (or onto) the human body, is another trend in organ printing. Inspired by fictional cinema, this technology could rapidly rehabilitate the tissues and recreate organs damaged by disease or on battlefields. This technology is no longer restricted to fiction. Currently, in situ cell printing has been tested to regenerate mouse skulls human external organs, such as skin

189.

188

and

With recent advancements in robot-

assisted surgery, cell printing will further lead the evolution of this technique to reality in the near future.

Author Contributions #These

authors contributed equally.

Funding sources This research was supported by the Bio & Medical Technology Development Program of the National Research Foundation (NRF) funded by the Ministry of


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Science & ICT (2017M3A9C6032067), the ICT Consilience Creative Program (IITPR0346-16-1007), the Government of Korea (MSIP) (2010-0018294), and the Korea Institute of Planning and Evaluation for Technology in Food, Agriculture, Forestry and Fisheries (IPET) through the Agri-Bio Industry Technology Development Program of the Ministry of Agriculture, Food and Rural Affairs (MAFRA) (316031-3). This research was also supported by the Institute for Information & Communications Technology Promotion (IITP) grant funded by the Korean Government (MSIT) (2017-0101982, Development of 3D printing-based artificial skin model for replacement of animal test and its commercialization). References 1. Hunsberger, J.; Neubert, J.; Wertheim, J. A.; Allickson, J.; Atala, A., Bioengineering priorities on a path to ending organ shortage. Current Stem Cell Reports 2016, 2 (2), 118127. 2. Tibbitt, M. W.; Rodell, C. B.; Burdick, J. A.; Anseth, K. S., Progress in material design for biomedical applications. Proceedings of the National Academy of Sciences 2015, 112 (47), 14444-14451. 3. Griffith, L. G.; Naughton, G., Tissue engineering--current challenges and expanding opportunities. Science 2002, 295 (5557), 1009-1014. 4. Khademhosseini, A.; Langer, R., A decade of progress in tissue engineering. Nature protocols 2016, 11 (10), 1775. 5. Park, J. Y.; Gao, G.; Jang, J.; Cho, D.-W., 3D printed structures for delivery of biomolecules and cells: tissue repair and regeneration. Journal of Materials Chemistry B 2016, 4 (47), 7521-7539. 6. Zhang, Y. S.; Yue, K.; Aleman, J.; Mollazadeh-Moghaddam, K.; Bakht, S. M.; Yang, J.; Jia, W.; Dell’Erba, V.; Assawes, P.; Shin, S. R., 3D bioprinting for tissue and organ fabrication. Annals of biomedical engineering 2017, 45 (1), 148-163. 7. Murphy, S. V.; Atala, A., 3D bioprinting of tissues and organs. Nature biotechnology 2014, 32 (8), 773. 8. Kengla, C.; Kidiyoor, A.; Murphy, S. V., Bioprinting Complex 3D Tissue and Organs. In Kidney Transplantation, Bioengineering and Regeneration, Elsevier: 2017; pp 957-971. ACS Paragon Plus Environment


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For Table of Contents Use Only Recent Strategies in Extrusion-based 3D Cell Printing toward Organ Biofabrication Ge Gao1, #, Byoung Soo Kim 1, #, Jinah Jang2, *, and Dong-Woo Cho 1, *



Figure 1. Currently available ECP systems: (A) pneumatic micro-extrusion including valve-free (left) and valve-based (right), (B) mechanical micro-extrusion including piston (left)- or screw (right)-driven microextrusion. Reproduced with permission from ref 12. Copyright Elsevier. 209x87mm (300 x 300 DPI)


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Figure 2. Optimized design of bioinks for reproducing the physical features of organs. (a) fabrication of a physiologically relevant geometry 3D printed structure using a silk-gelatin bioink (upper: CT scan of patient, segmentation of cheek geometry, lower: printed cheek geometry) Reproduced with permission from ref 96. Copyright Elsevier; (b) centimeter size complex objects using a mixture of 10 % (w/v) bovine gelatin, 0.5 % (w/v) alginate, and 2 % (w/v) fibrinogen (left: honeycomb structure and centimeter-size complex objective with overhanging features, middle: 3D models for printing a human ear, right: printed adult size ear). Reproduced with permission from ref 98. Copyright Wiley-VCH; (c) alginate- CNCs bioink formulation to fabricate a liver-mimetic construct (left: bioink formulation, right: comparison of printed structures using alginate and alginate-CNC bioink). Reproduced with permission from ref 109. Copyright Elsevier; (d) interpenetrating network (IPN) bioinks for printing; IPN bioink that was synthesized by covalently crosslinking PEG and ionically crosslinking alginate and cyclic mechanical deformation. Reproduced with permission from ref 89. Copyright Wiley-VCH. 209x130mm (300 x 300 DPI)



Figure 3. Representative examples of evolved ECP strategies. (a) Schematic of a printing complex 3D structure in a supporting bath made from gelatin slurry followed by bath banishment, which results in the fabrication of a chicken heart model (from left to right in the bottom: 3D CAD model, fluorescence, and dark-field view, scale bar: 1 cm). Reproduced with permission from ref 128. Copyright American Assoc. Advancement Science; (b) schematic illustration revealing the advantage of in-situ-crosslinking in printing delicate structures compared with pre-/post-crosslinking methods. Reproduced with permission from ref 139. Copyright Wiley-VCH; (c) schematic of an alginate-based bioink printing supplemented with an aerosolspraying process, producing semi-crosslinked struts (orange: non-crosslinked, purple: crosslinked by exposure to CaCl2 aerosol). Reproduced with permission from ref 138. Copyright American Chemical Society; (d) conceptual diagram of thermal-controlled 3D cell printing applying bottom and top heating and the resultant fabrication of a liver-shaped construct (left: top view, right, side view). Reproduced with permission from ref 147. Copyright Nature Publishing Group; (e) a human-scale ear tissue construct fabricated through the use of a multiple nozzle equipped ITOP system. Reproduced with permission from ref 149. Copyright Nature Publishing Group; (f) Schematic illustration of coaxial nozzle enabled 3D cell printing, where the CaCl2 (blue dots) in the shell are simultaneously printed with the bioink (GelMA (red), alginate (green), photoinitiator, and cells (grey)) in the core, and induced gelation of the alginate component. Reproduced with permission from ref 150. Copyright Wiley-VCH. 209x145mm (300 x 300 DPI)


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Figure 4. Representative examples that show indirect and direct extrusion techniques for vascularization strategy. (a) the concept of omnidirectional printing of 3D microvascular networks within a hydrogel reservoir. Reproduced with permission from ref 155. Copyright Wiley-VCH; (b) 3D convoluted renal proximal tubule on chip; schematic of the indirect fabrication process using PF127 (left); a 3D rendering of the printed convoluted proximal tubule acquired by confocal microscope (right). Reproduced with permission from ref 156. Copyright Nature Publishing Group; (c) schematic diagram that describes indirect extrusion using agarose. Reproduced with permission from ref 154. Copyright Elsevier; (d) coaxial 3D cell printing; (i) a schematic illustration of vascular-like structures of the coaxial extrusion process; (ii) an image of bioprinted alginate microfluidic channels; (iii) cell viability of alginate vascular conduits. Reproduced with permission from ref 78. Copyright The American Society of Mechanical Engineers; (e) coaxial bioprinting of a 3D construct (10 layers stacking) with perfusable hollow tubes using a blended bioink composed of alginate, GelMA, and PEGTA, which permit the encapsulated vascular cells to achieve early maturation of vessel tissue on day 21 (green: α-SMA, red: CD31, blue: DAPI). Reproduced with permission from ref 158. Copyright Elsevier. 209x239mm (300 x 300 DPI)




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Figure 5. Representative examples that show tissue-specific bioinks and their applications. (a) decellularization of cartilage, heart, and adipose tissues; (i) the images of decellularized tissues and histological analysis compared with native tissues; (ii) biochemical analysis of developed bioinks; (iii) fabricated cell-laden constructs with dECM bioinks.; (b) evaluation of differentiation into tissue-specific lineages and structural maturation of cells; immunofluorescence images showing chondrogenic differentiation of hTMSC in (i) COL and (ii) cartilage dECM constructs presentation COL type II (red), cell nuclei (DAPI, blue), and F-actin (green) (scale bar: 50 um); structural maturation of myoblasts in (iii) COL and (iv) heart dECM constructs displaying Myh7 (red) and cell nuclei (DAPI, blue) (scale bar, 200 um); adipogenic differentiation of hASCs in (v) COL and (vi) adipose dECM showing PPARγ (red), COL IV(green), and cell nuclei (DAPI, blue) (scale bar, 50 um) . Reproduced with permission from ref 165. Copyright Nature Publishing Group; (c) Fabrication of dome-shaped construct (i); evaluation of adipogenic differentiation without exogenous factors with adipogenic markers (PPARγ (red), COL IV (green), and cell nuclei (blue)) in vitro (ii); H&E and Immunohistochemical images (cell nuclei, blue; collagen IV, green; PPARγ, red) of the implanted constructs at 12 weeks of post-implantation (iii), (iv) in turn. Reproduced with permission from ref 30. Copyright Elsevier; (d) Illustration of pre-vascularized stem cell patch including multiple cell-laden bioinks and supporting PCL polymer (i); visualization of the implanted patch at epicardium for each group (ii). Reproduced with permission from ref 153. Copyright Elsevier. 209x219mm (300 x 300 DPI)




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Figure 6. Representative examples of advanced bioink designs for providing dynamic environments to the printed constructs. (a) (up) schematic of fabrication a bilayer structure composed of two PEG with different swelling ratio using photocrosslinking steps, (down-left) a deconvoluted fluorescent image of self-folded bilayer with fibroblasts stained with hoechst (blue) and calcein AM (green) in the inner and out layer, respectively, and (down-left) the insulin production of β-TC-6 cells cultured within self-folded micropatterned hydrogel (blue bar: basal insulin secretion (5.5 mM glucose, red bar: stimulated insulin secretion (16.5 mM glucose)). Reproduced with permission from ref 174. Copyright Elsevier; (b) (up) schematic illustration of shear stress-induced orientation of cellulose fibers during 3D extrusion process and resultant effect on anisotropic stiffness E and swelling strain α, and (down) flower-shaped construct composed of 90°/0° (left) and -45°/45° (right) bilayers oriented with respect to the long axis of each petal, with time-lapse sequences of the flowers during the swelling process (bottom panel) (scale bars, 5 mm, inset = 2.5 mm). Reproduced with permission from ref 23. Copyright Nature Publishing Group. 209x241mm (300 x 300 DPI)


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