Wake Forest Institute for Regenerative Medicine (WFIRM) scientists have improved upon the 3D bioprinting technique they developed to engineer skeletal muscle as a potential therapy for replacing diseased or damaged muscle tissue, moving another step closer to someday being able to treat patients.
Skeletal muscles are attached to bones by tendons and are responsible for the body’s movement.
When they are damaged, there is often loss of muscle function because the nerves are no longer sending signals to the brain.
Treatment of extensive muscle defect injuries like those caused by improvised explosive devices (IEDs) on the battlefield, for instance, is difficult and often requires reconstructive surgery with muscle grafts.
Effective nerve integration of bioengineered skeletal muscle tissues has been a challenge.
“Being able to bioengineer implantable skeletal muscle constructs that mimics the native muscle to restore function represents a significant advance in treating these types of injuries,” said lead author Ji Hyun Kim, PhD, of WFIRM.
“Our hope is to develop a therapeutic option that will help heal injured patients and give them back as much function and normalcy as possible.”
The work is detailed in a paper published online today by the journal Nature Communications.
Institute scientists previously demonstrated that the Integrated Tissue and Organ Printing System (ITOP), developed in house over a 14-year period, can generate organized printed muscle tissue that is robust enough to maintain its structural characteristics.
Since then, the WFIRM researchers have been developing and testing different types of skeletal muscle tissue constructs to find the right combination of cells and materials to achieve functional muscle tissue.
In the current study, they investigated the effects of neural cell integration into the bioprinted muscle construct to accelerate functional muscle regeneration.
“These constructs were able to facilitate rapid nerve distribution and matured into organized muscle tissue that restored normal muscle weight and function in a pre-clinical model of muscle defect injury,” said Sang Jin Lee, PhD., co-senior author, also of WFIRM.
This ongoing line of research at WFIRM is supported and federally funded through the Armed Forces Institute of Regenerative Medicine.
The goal is to develop clinical therapies to treat wounded warriors that will also benefit the civilian population.
WFIRM 3D bioprinter prints muscle. The image is Wake Forest Institute for Regenerative Medicine/WFIRM.
“Continued improvements in 3D bioprinting techniques and materials are helping us advance in our quest to make replacement tissue for patients,” said co-senior author Anthony Atala, M.D., who is director of WFIRM.
Funding: The research was supported by the Army, Navy, National Institutes of Health, Air Force, Veterans Administration, and Health Affairs to support the Armed Forces Institute of Regenerative Medicine II effort under award W81XWH-14-2-0004. The work was, in part, supported by the Medical Technology Enterprise Consortium grant 2017-614-002.
Additional co-authors include: Ickhee Kim, Young-Joon Seol, In Kap Ko, and James J. Yoo, all of WFIRM. Some or all of the authors are inventors on patent rights related to this work owned by Wake Forest University Health Sciences. The patents, whose value may be affected by publication, have the potential to generate royalty income in which the inventors would share.
Recent advances in regenerative medicine have confirmed the potential to manufacture viable and effective tissue engineering 3D constructs comprising living cells for tissue repair and augmentation.
Cell printing has shown promising potential in cell patterning in a number of studies enabling stem cells to be precisely deposited as a blueprint for tissue regeneration guidance. Such manufacturing techniques, however, face a number of challenges including;
(i) post-printing cell damage,
(ii) proliferation impairment and,
(iii) poor or excessive final cell density deposition.
The use of hydrogels offers one approach to address these issues given the ability to tune these biomaterials and subsequent application as vectors capable of delivering cell populations and as extrusion pastes. While stem cell-laden hydrogel 3D constructs have been widely established in vitro, clinical relevance, evidenced by in vivo long-term efficacy and clinical application, remains to be demonstrated.
This review explores the central features of cell printing, cell-hydrogel properties and cell-biomaterial interactions together with the current advances and challenges in stem cell printing. A key focus is the translational hurdles to clinical application and how in vivo research can reshape and inform cell printing applications for an ageing population.
Cell printing – capacity and limitations. Cell type and density are important factors in the printing of living cells. Post-printing consequences including: (i) impaired cell proliferation, (ii) maintenance of phenotype and genotype, (iii) preservation of cell integrity and morphology are issues that need to be considered.
Stem cell printing aims to deposit cells in three dimensions within an environment conducive to proliferation and differentiation. Therefore, long-term investigations of cell viability and proliferation in vitro and in vivo are required to elucidate construct maturation and effective tissue regeneration and integration. Importantly, the properties of the hydrogel cell carrier (biocompatibility, bioactivity, physical characteristics) for the select 3D print approach envisaged are crucial for cell encapsulation, protection and support during differentiation and proliferation.
Cell printing offers significant promise as an engineering technology to orchestrate tissue and organ regeneration including application, at scale, for human tissue formation and reparation .
Since the early seminal studies using cell deposition (cytoscribing) with a common desktop inkjet printer , major advances have been achieved in the arena of cell printing. Cell printing intends to position gel-like materials loaded with living cells (referred to as bioink) in a layer-by-layer fashion through an automated dispensing system. The attraction lies in the ability to apply this tissue engineering (TE) technology to generate readily implantable, tissue-relevant, 3D structures  to enhance the reparative processes.
In contrast to the use of standard cell seeding approaches, cell printing technologies incorporate cells within 3D implants to provide an improved biomimetic cell/material arrangement. Currently, two-dimensional monolayer cell culture remains the conventional platform for cell expansion and in vitro investigations.
However, cells are naturally able to sense their surrounding three-dimensional environment resulting in adaptations to growth and differentiation . To address these limitations, cell printing technologies have come to the fore, through the application of computer-based motion controls and biomaterial cell-carriers to fabricate 3D bioconstructs. These cell-laden scaffolds can recapitulate the 3D cell-niche environment including spatially organised homing signals essentials for tissue regeneration purposes.
This innovative tissue engineering technology, whilst widely accepted, presents a number of challenges to in vivo efficacy for tissue fabrication. These include:
- i)the requirement for biomaterials used for cell printing to mimic specific tissue extracellular matrix (ECM) physical and chemical properties,
- ii)the requirement for viscoelastic properties to allow both post-printing stability and appropriate fluidity for cell protection within the printing nozzle [5,6], and
- iii)the ability to preserve the viability of embedded cells during printing and host functional viable cells post-printing until remodelling and regeneration is complete.
This review will focus on recent research on 3D printing and in particular the challenges associated with printing living cells to generate viable and functional three-dimensional tissue constructs.
We will describe the use of various biomaterials commonly applied for cell printing and their limitations with a particular focus on those of promise for skeletal regeneration. Finally, we discuss the current challenges in cell printing toor engineer living tissues, listing relevant studies on how cell density, shear stress, printing parameters, nozzle shape and crosslinking, can affect post-printing viability and functionality.
We will highlight, in particular, current advances and challenges in skeletal stem cell printing including the translational hurdles that need to be overcome to ensure clinical application and how n vivo research can reshape and inform cell printing for therapeutic application.
Cell printing: state-of-the-art
Cell printing is developing apace with input and innovation from a range of disciplines including engineering, physics, materials chemistry and biology. Developments include the novel use of 3D printing technology incorporating bacteria in a hydrogel to generate a “functional living ink” that can impact in bioremediation (e.g. phenol waste removal) and biomedical (e.g. in situ biomaterial production) applications .
Tissue engineering application of cell printing are many but typically, focus on the manufacture of cell-laden 3D constructs to resemble the geometry and complexity of human tissues. Harnessing the biofabrication rationale (Fig. 1), cells are isolated from donor tissue, encapsulated within cytocompatible polymeric matrices and printed at a resolution that matches the heterogenic components of natural tissue in the 10–100 μm range .
Hydrogels are highly hydrated polymeric matrices  commonly used as biomaterial inks for 3D bioprinting. Three-dimensional hydrogel matrices (Fig. 2) can function as an injury site ECM scaffold for stem cell-mediated tissue-regeneration, or to deliver bioactive molecules to promote endogenous progenitor cell migration and differentiation .
Hydrogels have been engineered for application as cell carriers for a plethora of cell printing systems. In the following sections we describe the key cell printing technologies in current use, including inkjet printing followed by laser  and extrusion-based bioprinting [12,13], and characterise a number of hydrogels that can be utilised as structural support for cells.
Inkjet-based bioprinting, first developed by Thomas Boland (Clemson University) in 2003 [14,15], is a widely employed, low cost, high speed 3D printing technology. 3D inkjet printing platforms were first optimised for generating 3D constructs from a commercial 2D ink-based desktop printer in 2008 [, , , ].
Liquid state hydrogels were deposited as a defined spot using thermal-, acoustic- or electromagnetic-induced physical displacement. The volume of the droplet can vary between 1 and 100 pl which equals to 1–30 μm in size . The thermal or mechanical stresses employed for extrusion purposes represent a considerable disadvantage for this type of printer with cell density and cell death during printing significant limitations.
Droplets low in cell percentage can be controlled with high efficiency and thus, to some extent, reduce shear stress and machine nozzle blockage. However, low cell density affects viability and, importantly, the tissue formation capability of the constructs. Modern inkjet printers are able to handle hydrogels in their liquid state with low viscosity given their droplet/jet physical formation.
In a recent study, Gao et al. printed human mesenchymal stromal cells (MSCs) in a layer-by-layer fashion [19,20]. Thus, poly(ethylene glycol) alone  or dimethacrylate – gelatin methacrylate (PEG-GelMA)  mixed with MSCs were printed into a cylindrical construct and cultured for up to 21 days stimulating early differentiation into bone and cartilage resembling tissues. Difficulties in the specific selection of bioink viscosity were reduced by modifying the mechanical properties of materials such as PEG-GelMA.
In this way, the authors printed cell-laden scaffolds to achieve an improved uniform distribution of MSCs within a deposited matrix in contrast to non-printed PEG-GelMA scaffolds. While low viscosity ensures good cell viability during printing, this presents challenges for shape fidelity. A recent approach applied combinations of several highly printable bioinks (e.g. alginate-pluronic PE6800, alginate-PEG) using permanent and sacrificial ink to optimise the printing of complex shape constructs . The multi-head inkjet printing system in this case was able to deposit multiple materials, and by the inclusion of muscle progenitor cells in the sacrificial ink, to produce a perfusable 3D in vitro model.
An alternative approach to improving shape fidelity of low viscosity inks, is to harness rapid gelation properties upon deposition. For example, Hedegaard et al.  used inkjet printing to deposit peptide amphiphiles (PAs) which, in protein-rich solutions, display instantaneous self-assembly to generate complex microstructures which could be modulated by tuning the inkjet propulsion, the impact speed and the bath solutions.
Although versatile and easy-to-use, inkjet cell printing technology is still, to date, not applicable to the development of human-size bioconstructs given issues with droplet non-uniformity, poor cell density, frequent nozzle blockage and physical stresses on cells (e.g. thermal and frequency shocks).
Laser bioprinting [, , , , ] is an effective process in preserving high cell viability following cell-deposition. This nozzle-free technique is based on laser-induced forward transfer (LIFT) physics through which cells seeded on a donor slide (covered in a radiation absorbing layer) can be safely propelled, encapsulated within droplets of biomaterial, toward a collector slide. The bioconstruct droplet resolution is influenced by the biomaterial rheological properties, donor-collector system, laser energy and resolving power. Laser bioprinting approaches require specific biomaterials with defined viscoelastic properties to ensure that the biomaterial/cells constructs can rapidly gelate to ensure high-defined shape retention .
LIFT-based bioprinting have been reported to offer the most useful approach for two-dimensional patterning of cells, yet a few recent studies have shown the potential to produce complex 3D patterns using this approach. Xiong et al.  applied matrix-assisted pulsed-laser evaporation direct-write (MAPLE DW) approaches to print 3D alginate-fibroblast hollow tubes. Despite the development of such a printing approach to enable a highly accurate and detailed 3D construct, the resultant printing process was slow.
In addition, cell viability, noted to be 63.8% immediately upon printing was observed to be 68.2% after 24 h. Low cell survival rate could be caused by both cell injury upon printing (mechanical stress during droplet/jet formation and landing) and the requisite stationary conditions (45 min) to allow gel crosslinking.
Using a similar approach, Gruene et al.  successfully generated a MSCs graft using LIFT technology with evidence of chondrogenic and osteogenic differentiation evidenced by osteocalcin (OC) and alkaline phosphatase (ALP) activity. Hence, stem cells were able to survive the print process and retain their osteogenic differentiation capacity.
Technologies such as stereolithography (SLA) or digital light processing (DLP) are reliable alternatives to LIFT for producing viable cell-laden constructs. SLA and DLP systems utilise the photopolymeric properties of the material (or compounds mixed with it) to solidify the liquid and thus form a 3D construct.
The curing process is commonly carried out simultaneous with the printing via exposure to ultraviolet (UV)  or visible light . Recently, Lim et al.  have employed methacrylated poly(vinyl alcohol) (PVA-MA) and gelatin (Gel-MA) to fabricate cell-laden constructs via direct visible light processing proving the high resolution (25 μm) printing platform is suitable for producing viable and functional cell-laden constructs.
At present, laser-based techniques offer the possibility to print cells with low damage, although, critically, require (i) specific bioink gelation properties for application, (ii) expensive fabrication costs and, (iii) a skilled workforce to operate the platform technology.
Extrusion-based bioprinting is a widely used rapid prototyping approach, able to deposit precisely hydrogels with shape retention depending on the physical and chemical properties of the biopolymer used. Extrusion-based bioprinting is currently one of the most widely employed platforms for cell printing given the advantages of ease of handling, ability to customise and versatility of the systems available [, , ].
Extrusion-based bioprinting typically uses post-processing or temperature-sensitive hydrogels for cell delivery. Post-printing processes (e.g. cross-linking, UV-curing, etc.) are widely employed for altering physio-chemical gel properties accelerating gelation phenomenon, strengthening the overall matrix structure and for tuning polymeric degradation. The use of specific hydrogels as cell carriers can highly influence cell viability after printing.
Fedorovich et al.  investigated the differentiation potential of MSCs within organised cell-laden constructs. Cells were included within different types of hydrogels such as synthetic Lutrol F127 (PF127), matrigel, alginate and agarose. Lutrol and agarose extrusion resulted in precise deposition with a 150 μm nozzle tip, however with limited cell survival following printing in comparison to matrigel and alginate, emphasising the influence of material choice in the cell printing process.
Cells can be encapsulated in modified-composite bioinks that enable shape fidelity post-printing even at low polymeric concentrations, preserving cell viability and sustaining proliferation. We have recently  demonstrated the ability of a novel clay-based hydrogel to encapsulate human skeletal stem cells and to generate high resolution cell-laden 3D scaffolds. The nanocomposite bioink was developed using a nanosilicate suspension blended with alginate and methylcellulose to ameliorate shear thinning properties, shape fidelity, mechanical stability and growth factor localisation.
Following extrusion, 70–75% of the printed cells survived the process, showing augmented proliferation rate compared to the control group (clay-free). This low-polymer content bioink facilitated cell printing, retaining viability after printing and supported cell proliferation over 21 days of culture.
Extrusion-based cell printing remains a popular approach given the wide versatility available in the system but, remains limited in resolution [, , ]. Cell printing, based on extrusion techniques, represent an exciting approach for organ printing given, i) the plethora and diversity of compatible materials that can be used, ii) the potential for cell inclusion and, iii) parameter variation available as a consequence of further modulation of printer parameters and scalability.Go to:Hydrogels for cell printing
3D printed scaffolds for tissue regeneration purposes can involve biomaterials alone (cell-free), cells alone (scaffold-free) or a combination of the two (cell-laden). The majority of biomaterials used in cell printing approaches are hydrogels. The high water content of hydrogels can promote nutrient diffusion and waste removal, while, manipulation of hydrogel properties can enable the rapid or prolonged release of a drug or molecule of interest [10,37]. To be employed as a cell-encapsulation vector for 3D printing applications, hydrogels should present specific rheological and gelation properties tuned according to the 3D fabrication process. The ultimate purpose for such polymeric matrices is to direct and guide tissue-specific cell lineage formation and to maintain cell proliferation and phenotype .
Cell-free, scaffold-free and cell-laden hydrogel printing
Cell-free (acellular) biomaterial deposition processes, with further post-printing cell addition, remain the most common bioprinting techniques in use. Although printing viscous hydrogels can enable the fine deposition of the material with subsequent cell-seeding, results from acellular approaches are far from ideal.
The acellular approaches may lack the requisite functionality critical for in vivo regeneration. Nevertheless, acellular implants are frequently used in the clinic for bone repair. Recently, Reznikov et al.  used a 3D printing approach to generate acellular 3D scaffolds incorporating a stiffness gradient which was able to modulate the response of endogenous progenitor cells in a large animal defect model (Fig. 3 Acellular). Acellular implants have been reported to show substantial tissue ingrowth in vivo but with a lower degree of ECM organization compared to cell-laden 3D hydrogel implants .
It is important to note, post-printing seeding approaches cannot ensure the requisite cell density, cell spreading, attachment, migration and interaction both in vitro and in vivo. Acellular 3D printed constructs thus have limited capacity for cell homing, restricting space available for post-seeded cells. Nevertheless, hydrogel coating  or hydrogel-mediated cell seeding  could represent a promising approach especially in hard tissue regeneration applications. For example, 3D printed cell-free ceramic such as β-tri-calcium phosphate (β-TCP) can be perfused with softer cell-laden hydrogels for effective cell delivery and encapsulation [9,43].
Scaffold-free approaches have come to the fore, as promising manufacturing techniques, to engineer 3D constructs using, exclusively, living cells and the absence of any support or encapsulating biomaterial [23,, , ]. For example, Arai et al.  recently fabricated a scaffold-free tubular cardiac implant and proved the efficacy of this construct as a cardiac pump. (Fig. 3 scaffold-free).
Cardiac spheroids were run through an array of fine needles using a computer-controlled robotic arm. Spheroid fusion and the contraction of the three-dimensional construct implied the potential functionality of this construct. Taniguchi et al.  used an unconventional Regenova bioprinter to build a fully functional tracheal substitute via Cyfuse’s method.
In a similar approach to Arai et al.  functional multi-cellular organoids (chondrocytes, endothelial cells and mesenchymal stem cells) have also been deposited and fused into a tubular shape. The results of this organoid printing approach, though promising when constructs were implanted in vivo, raised questions of practical feasibility. Issues included: i) the high number of spheroids required (384 spheroids for 5 mm long construct), ii) poor structural integrity (artificial trachea resulted significantly less strong than 8-week old rat trachea) and, iii) the long timescale required for full implant maturation (28 days).
Cell-laden printing approaches can manufacture biomimetic human-like tissues with stem cells encapsulated in biomaterials able to initiate and stimulate new tissue ingrowth acting as primitive building blocks (Fig. 3 cell-laden) [30,31,33].
The cell-carrier hydrogels are typically doped with GFs, bioactive compounds or macromolecules to aid cell metabolism [9,47]. The hydrogel design should therefore consider both cell microenvironment and printing capabilities. High printing accuracy can be achieved due to tuned gel viscosity reducing strand failure.
Nevertheless, the selected biomaterial for cell encapsulation should retain the capacity to be extruded with minimum applied shear force to avoid cell damage and reduce cell viability [48,49]. Biomaterials commonly used in combination with cell populations for cell-laden printing include natural, synthetic and hybrid (combination of natural and synthetic) bioinks. The desired material selected according to the functional ultimate use, to try to circumvent inherent material limitations for the printing of living cells.
Hydrogels for cell printing – ECM matrix biomimetics
Biomaterials selected for cell printing should, ideally, allow the cells to regain their original shape following extrusion through a fine aperture and provide the cells with the appropriate stimuli to induce proliferation and differentiation post printing and preserve cell spatial location [40,41]. Hydrogels engineered to recapitulate the ECM environment, are typically composed of polymeric chains that are often crosslinked with one another [51,52]. Such ECM resembling hydrogels for cell printing derived from decellularised ECM (dECM) cartilage (cdECM), adipose (adECM) and heart tissues (hdECM) have been investigated by Pati et al. . Human adipose-derived stem cells (hASCs) and human inferior turbinate-tissue derived mesenchymal stromal cells (hTMSCs) were used as a model cell line to investigate the inherent phenotype tissue-mimicking gel properties. DECM provided a temperature-sensitive hydrogel gelling at physiological temperatures with >95% cell viability at 24 h. HdECM  was used to develop a viable 3D printed patch able to induce a potent angiogenic response and enhance cardiac function in vivo.
Fedorovich et al.  embedded MSCs within a matrigel comprising calcium phosphate (CaP) nanoparticles and investigated their efficacy for bone formation in a subcutaneous mouse model. Calcium phosphate components have been shown to be particularly beneficial for new bone formation both in vitro and in vivo through the release of calcium and phosphorus ions emphasising the attraction of such an approach. The size requisites to enable extrusion-based systems, typically, force the use of material particles in the nanometres range. In vitro investigations demonstrated high MSCs viability and bone formation after 6 weeks, although, only fibrous-like tissue was observed within in vivo implants showing lack of the desirable osteoinductive properties.
Hydrogels – limitations in cell printing applications
Bioinks lack inbuilt mechanical support
Cell printing can produce stable implants but these typically, lack any tissue specific mechanical properties. Support materials are often used to provide further mechanical aid to the final 3D printed tissue substitute. To address this issue, a number of groups have employed a hybrid approach, that is. the simultaneous deposition of a cell-laden hydrogel with a supportive material to generate a mechanically competent 3D printed constructs [1,, , , , ].
Poly(ε-caprolactone) (PCL) is a synthetic polymer capable of withstanding a wide range of external loads in comparison to soft and high water-content gels, although, the requisite high extrusion temperature generally makes PCL unsuitable for cell printing. Even as a mechanically stable material, the hard nature of the PCL surface could result in disruption to soft hydrogels providing a non-homogenous environment for cell expansion (altered cell-surface interaction modulating cell proliferation) [1,50].
While hydrogel viscosity can be tuned and can increase linearly with polymer concentration, care is often required to ensure a cell compatible environment  and allow for degradation within a relevant time-frame following implantation. Recently, Freeman et al.  defined a printability window by looking at differences in alginate molecular weight (MW) in combination with modulation of ionic crosslinking addition. Varying MW to crosslinking ratio, degradation and encapsulated growth factor release could be tuned according to the specific need of developing a hard or soft region.
A further approach to overcome the issue of printing soft hydrogels is the application of a gel bath [, , ]. The temporary supportive gel allows aiding printing of low polymeric bioinks, building larger and more complex structures. After printing, the constructs can be crosslinked to enhance further the mechanical properties of the scaffold.
Low polymeric content hydrogels are ideal for cell printing
Hydrogel stiffness, typically correlated with polymer percentage content, can directly influence cell survival and proliferation rates post-printing. High polymer content hydrogels used for printing can impose higher stress on encapsulated cells than softer hydrogels. Physical stress applied on cells by a stiff hydrogel can prove detrimental for post-print cell viability.
Therefore, the optimal biomaterial for cell printing should provide a balance in polymer content ensuring a soft and porous three-dimensional microenvironment where cells can be encapsulated avoiding excessive physical stresses as with high polymer materials [31,, , ]. Hydrogel physical properties such as high polymer content can influence cell survival by hindering nutrient supply and waste removal. Indeed, scaffold porosity is recognised as central to in vivo environments to elicit appropriate oxygenation and blood vessel ingrowth. It is well known that scaffold pore sizes can influence rate of diffusion and certain pore sizes) may be optimal for, variously, revascularization (5 μm), skin (20–125 μm) and bone (100–400 μm) regeneration .
The specific spatial distance between cells encapsulated within deposited fibres and outer oxygen rich ambient can influence nutrient and waste diffusion however, the very nature of hydrogels, being predominantly water, can generally facilitate this exchange, supporting cells during their proliferation process.
Natural or synthetic hydrogels can influence printability and cell behaviour
Natural materials, derived from natural polymers, often convey the added advantage of cytocompatibility and the provision of natural cues to provide a favourable microenvironment for cell differentiation [30,33]. Synthetic biomaterials synthesized from polymers or blocks of co-polymers, typically yield more consistent and reproducible structures than natural polymers. Synthetic and natural biomaterials have however been employed as surfactant agents, bio-carriers and support material for several cell lines  and can be applied for cell printing application given their viscoelastic mechanical properties and shear thinning behaviour [31,60]. Additionally, hydrogels selected for bioprinting should demonstrate stable gelation conditions to enable even deposition and ensure shape fidelity [67,68]. Hybrid materials can be applied to create optimise inks for encapsulating and printing living cells as they combine the natural and the synthetic components to produce highly printable bioinks [31,69,70].