Posts Tagged ‘scaffold’

First 3D Printed Tibia Replacement

Reporter: Irina Robu, PhD

Current advances have allowed 3D printing of biocompatible materials, cells and supporting components into complex 3D functional living tissues. 3D bioprinting has already been used for the generation and transplantation of several tissues, including multilayered skin, bone, vascular grafts, tracheal splints, heart tissue and cartilaginous structures. Thanks to 3D printing, an Australian man got to keep his leg. The man, Reuben Lichter nearly lost his leg above the knee due to a bacterial infection. Doctors told him that he had osteomyelitis which infected his entire bone. Lichter’s bacterial disease of osteomyelitis affects 2 in every 10,000 people in the United States. He had two choices: an experimental procedure using the 3D printed bone or lose his leg. For Lichter, the choice was easy.

Michael Wagels who served as the lead surgeon performed the world’s first-ever transplant surgery using a 3D printed bone. The scaffold was initially modeled at Queensland University of Technology. Biomedical engineers designed the scaffold to promote bone growth around it and then slowly dissolve over time. To have the body successfully grow around the scaffold, the team introduced tissue and blood vessels from both of Lichter’s legs to the scaffold. The surgery itself happened over five operations at Brisbane’s Princess Alexandra Hospital.

However, the next major challenge for biomedical engineers is how to successfully 3D print organs.



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Platform Technologies for Directly Reconstructing 3D Living Biomaterials

Reporter: Irina Robu, PhD

The techniques of electrospraying and electrospinning have existed for at least a century. These techniques employs a high voltage applied to a needle accommodating the flow of media, placed above a counter electrode which could either be grounded or have an opposite charge to the needle—thus introducing the charged media to an electric field.

These endeavors have demonstrated the wider applicability of these technologies and hence in the last 20 years or so have been used for the direct handling of a wide range of materials, including bio-inspired materials. These investigations have generated interest in areas such as the development of fine monolayered surfaces, fabrication of scaffolds which could be used for many laboratory-based fundamental biological studies.

In 2005, Jayasinghe et al. began investigations into both electrospraying and electrospinning of immortalized cell lines. Even though the high voltages involved, these cells were  found to be viable post-electrospraying/electrospinning. Additional work has extended these studies to different cell types, both murine and human, immortalized or primary, stem cells, and even whole fertilized embryos from model organisms. Established protocols (such as flow cytometry, genetic/genomic interrogation, and microarray analysis) proved that cells processed using either electrospraying or electrospinning were indistinguishable from controls. Hence bio-electrospraying (BES) and cell electrospinning (CE) have become platform technologies for the biological and life science and are the leading technologies for the direct handling of cells—both for distribution of cells with pinpoint precision as cell-bearing droplets, and for the formation of truly 3D living scaffolds.

Previous studies have been carried out with processed cells suspended in matrices generated from animal/tumor-derived materials which contain largely uncharacterized growth factors and bioactive signals. This makes them very undesirable for clinical assays. While not applicable to humans, they can be used  with advanced biopolymers, which could be directly translated to humans, and have the potential for creating artificial constructs which could be used for a variety of applications in the regenerative medicine field. The present study describes the in vivo application of such biopolymers, using murine macrophages to interrogate biocompatibility and cellular behavior post-transfer.



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3D BioPrinted Carbon Nanotubes used to Stimulate Bone Regrowth

Reporter: Irina Robu, PhD

Bone disorders are of significant concern due to increase in the median age of our population and at this present time bone grafts have are used to restore damaged bone. However, synthetic biomaterials are now being used as bone graft substitutes and they are selected for structural restoration based on their biomechanical properties. Lately, scaffolds are engineered to be bioactive to enhance tissue growth. These scaffolds are usually porous, made of biodegradable factors, drugs or stem cells.

The research group led by Dr. Maria Vallet-Regi at Faculty of Pharmacy-Universidad Complutense de Madrid showed that carbon nanotubes to the mix to create 3D electrical network within the bone tissue can stimulate bone cell regrowth. The polymer they used was polycarpolactone (PCL), which is rather easy to 3D print. According to Mercedes Vila, the Principal Investigator in charge of the project, the carbon nanotubes were added to the bio-printable material mixture to create a three-dimensional electrical conducting network all through the volume of the scaffold, which would allow the application of this stimulation to the scaffold once implanted on the damaged bone site.

“In this sense, electrical stimulation has been explored since the discovery of the presence of electrical potentials in mechanically loaded bones,” Mercedes pointed out. “Certain types of cell behavior, such as adhesion and differentiation, can be affected by the application of electrical stimulation. Thus, the creation of a permanent charge on the material surface, positive or negative, as well as a direct electrical stimulation can promote the attraction of charged ions from the environment to the cells. This would modify their protein adsorption with the subsequent influence on the cells’ metabolic activity. Therefore, the use of electrical stimulation after biomaterial implantation to favor cell adhesion and differentiation and, consequently, induce bone healing seems a smart approach to accelerate the osteointegration process.”

Adding CNTs into the bio-printed polymer and mineral prosthetic bone can stimulate regrowth of the actual bone cells. However, bio-printing CNTs created no extra difficulties, as they are so thin that they can be extruded with ease through any pneumatic syringe. Most of the complications are related to finding the correct viscosity in the combination of CPL and hydroxypatite.

“Finding the right right viscosity to be extruded through the syringe while keeping enough robustness to get the 3D scaffold printed at room temperature, was complicated,” Mercedes admitted. “At the same time as the slurry was prepared in dichloromethane solution for diluting the PCL, achieving the right viscosity while evaporating the solvent was tricky. Moreover, once the PCL and the hydroxyapatite were mixed together, the addition of the CNTs was performed and reaching a proper dispersion took a bit of stirring time.”

Using EnvisionTEC’s 3D bioplotter, the researchers were able to create very complex 3D structures which would enhance the future for tissue replacements as it allows tailored solutions by capturing the anatomical information of the patient’s wound by computed tomography and magnetic resonance, for example, to obtain a personalized and unique implant.

As with many other 3D printing applications, it appears we are only starting to scratch the surface of the possibilities that are ahead for bioprinting.



Other related articles published in this Open Access Online Scientific Journal include the following:

search for Bone related articles, please place here references

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Biofabrication with Stem Cells

Curator: Larry H. Bernstein, MD, FCAP



Biofabrication  Special Issue:  Dec 2015; 7(4).    http://iopscience.iop.org/1758-5090/7/4


Three-dimensional bioprinting of embryonic stem cells directs highly uniform embryoid body formation

Liliang Ouyang1,2,6, Rui Yao1,2,6, Shuangshuang Mao1,2, Xi Chen3, Jie Na3 and Wei Sun1,2,4,5
Biofabrication, Volume 7(4)    http://iopscience.iop.org/article/10.1088/1758-5090/7/4/044101/meta

With the ability to manipulate cells temporarily and spatially into three-dimensional (3D) tissue-like construct, 3D bioprinting technology was used in many studies to facilitate the recreation of complex cell niche and/or to better understand the regulation of stem cell proliferation and differentiation by cellular microenvironment factors. Embryonic stem cells (ESCs) have the capacity to differentiate into any specialized cell type of the animal body, generally via the formation of embryoid body (EB), which mimics the early stages of embryogenesis. In this study, extrusion-based 3D bioprinting technology was utilized for biofabricating ESCs into 3D cell-laden construct. The influence of 3D printing parameters on ESC viability, proliferation, maintenance of pluripotency and the rule of EB formation was systematically studied in this work. Results demonstrated that ESCs were successfully printed with hydrogel into 3D macroporous construct. Upon process optimization, about 90% ESCs remained alive after the process of bioprinting and cell-laden construct formation. ESCs continued proliferating into spheroid EBs in the hydrogel construct, while retaining the protein expression and gene expression of pluripotent markers, like octamer binding transcription factor 4, stage specific embryonic antigen 1 and Nanog. In this novel technology, EBs were formed through cell proliferation instead of aggregation, and the quantity of EBs was tuned by the initial cell density in the 3D bioprinting process. This study introduces the 3D bioprinting of ESCs into a 3D cell-laden hydrogel construct for the first time and showed the production of uniform, pluripotent, high-throughput and size-controllable EBs, which indicated strong potential in ESC large scale expansion, stem cell regulation and fabrication of tissue-like structure and drug screening studies.

With the capability of self-renewal and differentiating into all somatic cell types, embryonic stem cells (ESCs) hold great promise as an in vitro model system for studies in early embryonic development, as well as a robust cell source for applications in diagnostics, therapeutics, and drug screening [1]. Derived from the inner cell mass of a blastocyst, ESCs requires delicate culture condition and trend to cluster together, and in particular, forms three-dimensional (3D) cellular spheroids termed embryoid body (EB) [2]. In order to better understand stem cell niche and regulation of ESC differentiation and reprogramming, in vitro recapitulation of the spatial distribution of cells, cell–cell and cell–matrix interactions, is of paramount importance [35]. Compared with 2D monolayer culture, 3D cell culture is believed to confer a higher degree of clinical and biological relevance to in vitro model [6, 7], since the spatial arrangement of cells and extra-cellular matrix could influence cell differentiation and function both in vivo [8] and in vitro[9]. Therefore, reconstruction of 3D cell microenvironment is critical to directing stem cell fate and generating cell sources for tissue engineering, regenerative medicine and drug screening studies.

By mimicking some of the spatial and temporal aspects of in vivo development, EB is a basic 3D model for ESCs culture and differentiation studies. It was reported that the size and uniformity of EBs could vastly influence stem cell fate [1012]. Various methods have been used to fabricate such cellular spheroid, basically including static suspension, hanging-drop and multiwell culture, most of which doesn’t involve biomaterials. Static suspension method inoculate suspension of ESCs onto non-adhesive plate to allow cells spontaneously aggregate into spheroid. This method is easy to operate, but showed limited control over the EBs size and shape due to the probability that ESCs encounter each other accidentally [13]. Hanging-drop is a common method to produce size-controlled homogeneous EBs, where droplets of ESCs suspension are pipetted onto the lid of a Petri dish and EBs was generate by gravity after overturning the dish [14]. However, manual pipetting is labor intensive and the reproducibility varies with operators. Multiwell culture offers high-throughput solution for EB formation through cell aggregation in uniformly shaped microwell arrays but requires expensive microwell culture plates [10, 15]. Besides, there are few customized microwell culture plates available in the market.

Recent advances in bioprinting technologies facilitated the precise deposition of ESCs in a reproducible manner. Xu et al [16] and Shu et al [17] printed ESCs suspension solution into 2D patterns as hanging-drop approach for EB formation, without the cell-biomaterial interaction. Corr and Xie [18, 19] applied laser direct-write method in bioprinting of mouse ESCs together with gelatin. ESCs maintained the pluripotency while proliferation and formed EB. EB size can be controlled by cell density and colony size. However, these studies just generated 2D cellular array without 3D cell–matrix interactions, and cell–cell interaction happens within one drop but not among different drops. To better recapitulate the characteristics of in vivo cell microenvironment, 3D customized cell/matrix construct with macro-porous structure might be a preferred choice. To our knowledge, there has been no report about bioprinting of ESCs into 3D cell-laden constructs.

The extrusion-based temperature-sensitive 3D bioprinting technology was developed in our lab and has been utilized for bioprinting of hepatocytes [20], adipose tissue-derived stem cells (ADSCs) [21], C2C12 cells [22], hela cells [23] and 293FT cells [24]. Most commonly used biomaterials for this technology are gelatin and alginate. Gelatin, a type of denatured collagen, is widely used as a coating for feeder layer-free mouse ES cell culture. Alginate, extracted from brown algae, is proving to have a wide applicability in tissue engineering and drug delivery and also used in embedding mouse ESCs for EB formation [25]. It has been proved in many studies that encapsulation of ESCs in hydrogels would direct EB formation with the maintenance of pluripotency [2628]. Hence, we hypothesized that the bioprinting of 3D ESC-laden construct would maintain the stem cell pluripotency and address the challenges associated with the current methods for EB formation.

In this study, we investigated the feasibility of applying extrusion-based temperature-sensitive 3D bioprinting technology in bioprinting of ESCs with hydrogels into 3D macro-porous structure, with the maintenance of viability, pluripotency, cell growth and to direct EB formation. Printing process parameters were optimized to obtain a high cell survival rate (90%) after printing process and construct formation. Stem cell pluripotency was examined by the expression of stem cell markers (octamer binding transcription factor 4 (Oct4), stage specific embryonic antigen 1 (SSEA1) and a homeodomain-bearing transcriptional factor (Nanog)) and the ability to form EBs. The regulation of EB formation in the 3D bioprinted construct was systematically compared with commonly used methodology, where EB formation relies on cell aggregating as well as cell proliferation. Results demonstrated that this novel technology generated pluripotent, high-throughput, highly uniform and size controllable EBs under static culture condition without complex equipment. This study established the feasibility of fabricating 3D in vitro tissue-like model using ESCs for the first time, creating engineered microenvironment for pluripotent stem cells with the ability of placing cells and materials spatially in a reproducible manner.



3.1. 3D bioprinting and cell viability optimization

In this study, many process parameters, e.g. nozzle inner diameter, nozzle insulation temperature and chamber temperature were examined to optimize cell viability after 3D construct fabrication. It was demonstrated that larger nozzle diameter resulted in higher cell viability (figure 2(A)). Specially, the cell viability under Nozzle-160 μm (81.59% ± 1.74%) was lower than those under Nozzle-260 μm (88.06% ± 1.98%), Nozzle-410 μm (89.59% ± 0.71%) and Nozzle-510 μm (90.84% ± 1.02%), with significant differences. Nozzle diameter of 260 μm, 410 μm and 510 μm showed no significant differences in terms of cell viability.

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Figure 2. The influence of bioprinting parameters on ESC viability is determined by fluorescence live/dead staining. (A) The influence of printing nozzle inner diameter on ESC viability (Insu-30 °C and Cham-10 °C). (B) The influence of nozzle insulation temperature and chamber temperature on ESC viability. Insu-25 °C means keeping the nozzle insulation temperature at 25 °C. Cham-4 °C means setting the chamber temperature at 4 °C, and so as others. (C) The fluorescent staining images show the live (green) and dead (red) cells at different days during culture period. Scale bar: 100 μm.

Insulation and chamber temperatures were altered to study their influences on cell viability (figure 2(B)). As a positive control, ESCs/hydrogel mixture without bioprinting were stained with fluorescence live/dead reagent, and showed 93.14% ± 1.31% cell viability. When insulation temperature was set at 25 °C (labeled as ‘Insu-25 °C’), cell viability increased with the chamber temperature from 55.52% ± 2.37% under 4 °C (labeled as ‘Cham-4 °C’) to 78.22% ± 2.55% under 10 °C (labeled as ‘Cham-10 °C’) with significant differences. When the insulation temperature was set at 30 °C (labeled as ‘Insu-30 °C’), nearly 90% ESCs remained alive under the chamber temperature of 7 °C and 10 °C (labeled as ‘Cham-7 °C’ and ‘Cham-10 °C’), significantly more than that under Cham-4 °C (72.40% ± 2.46%). To achieve both high ESC viability and a clear construct configuration, the process parameter combination of Nozzle-260 μm, Cham-10 °C and Insu-30 °C was chosen.

After culturing for three days, few cells were found dead, which were isolated from living EBs (figure 2(C)). On day 5 and day 7, a few dead cells were observed on the edge of EBs. About 5% ESCs were stained dead on day 7. As the static culturing continued, 9.69% ± 1.77%, 17.72% ± 2.91% and 40.64% ± 2.06% were found dead on day 8, day 9 and day 10, respectively (supplement 2). So, we chose 7 days as the culture period in the following analysis.

3.2. Construct structural stability and EB formation

A 3D cellular construct with the cross section of 8 mm × 8 mm and height of 1 mm was fabricated under the optimized process parameter. The 3D construct demonstrated macro-porous grid structure in which the hydrogel threads were evenly distributed according to the computer design (figure 3(A)). Both the width of the threads and the gap between the threads were homogeneous, that is 728.2 μm ± 24.9 μm and 424.3 μm ± 17.8 μm, respectively, suggesting 3D cellular construct formation in a highly controlled manner. ESCs were embedded uniformly in the hydrogel matrix threads, developing a specific 3D microenvironment.

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Figure 3. Images of the printed cellular model with grid structure. (A) Full view of the cellular construct. (B) Phase-contrast images demonstrating the cell morphology and distribution of different cell density at day 3, day 5 and day 7. Scale bar: 1 mm.

During the culture period, ESCs tended to grow as spheroid cellular aggregates, also known as EB. The cell density in the 3D hydrogel construct were determined by the initial cell density in the ESC/alginate/gelatin mixture and showed significant influence on the yield and density of EBs formed in the construct (figure 3(B)). It was demonstrated by semi-quantitative analysis of figure 3(B) that, the percentage of area occupied by EBs varied from 52% to 85% when initial cell density changed from 0.5 mln mL−1 to 2.0 mln mL−1 . Most of the EBs were contained in the hydrogel threads in the culturing period. However, when the initial cell density was as high as 2.0 mln ml−1, some of the EBs were observed running off from the threads into the throughout holes.

3.3. Cell proliferation

ESCs formed spheroid EBs in the 3D hydrogel construct and the diameter of the EBs enlarged with culturing time while keeping their spatial location in the hydrogel thread, indicating EB formation by ESC proliferation rather than aggregation (figure 4(A)). Compared with traditional 2D culture, ESCs showed different proliferation rate indicated by the OD value measure by CCK-8 kit (figure 4(B)). The normalized OD value of the 3D in situ group grew faster than that of 2D from day 1 to day 3, while slowing down after day 3 and being much less than that of 2D at day 7. However, 3D harvest group showed a generally faster growth rate than 2D during the one week culturing, with a significant difference. In addition, the diameter of EB was also measured to indicate ESC proliferation rate. When comparing the normalized EB volume with normalized 2D OD value, 3D samples also maintained a significantly faster growth rate than 2D, though the EB volume had huge variance (figure 4(B)).

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Figure 4. EB growing and cell proliferation. (A) Magnified images of the same location in 3D printed cellular construct at different times. (B) ESC proliferation in the 3D construct compared with 2D culture. All the date were normalized to the value of day 1. Scale bar: 200 μm.

Pluripotency markers, i.e. Oct4, SSEA1 and Nanog were analyzed to determine the pluripotency maintenance of ESCs after 7 day culture in the 3D hydrogel construct. Immunofluorescence staining and flow cytometry analysis showed that almost all of the cells within the EB were successfully stained both Oct4 and SSEA1. Because of the limitation of confocal capacity when dealing with large scale aggregates, the central part of the EB was darker than the edge (figure5(A)). Flow cytometry analysis demonstrated that 97.2% and 99.0% cells were positively stained with Oct4 and SSEA1 respectively (figure 5(B)). The qRT-PCR results demonstrated that the gene expression level of Oct4 and Nanog in our 3D samples were close to those in 2D (within the deviation of ±3%), without significant difference, confirming that cells have maintained pluripotency (figure 5(C)).

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Figure 5. ESC pluripotency at day 7 was determined by CLSM, flow cytometry and qRT-PCR. (A) Immunofluorescence images of EBs stained with Oct4, SSEA1 and DAPI. (B) Quantification of 3D dissociated cells marked with Oct4 and SSEA1 by using flow cytometry. (C) Gene expression of Oct4 and Nanog in 3D versus 2D by using qRT-PCR. Scale bar: 50 μm.

EBs were harvested from the 3D hydrogel construct at different time intervals to analyze EB morphology (figure 6(A)). Most of the EBs were separated without fusion. The center part of the EBs was darker than edge part, especially at day 5 and day 7, indicating the 3D sphere structure of EBs. Through analyzing the size of 250 random EBs for each sample, the histogram of EB diameter were obtained, showing a Gauss distribution curve (figure 6(B)). The results demonstrated that the EB size increased significantly from about 50 μm to about 110 μm when the construct was cultured from day 3 to day 7 (figure 6(C)). Cell density had little influence on EB average size. However, increased cell density would result in the reduction of the uniformity of EB size, especially at day 7; the EB diameter of 2.0 mln mL−1 group at day 7 was vastly heterogeneous, with a deviation of 42.30 μm, which was much more than those of other two groups.

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Figure 6. EB formation in different cell density: (A) optical images of released EBs at different days. (B) EB diameter and (C) EB circularity distributions at different days. Summary of the (D) diameter and (E) circularity. 250 EBs were applied for diameter and circularity measurements for each group. Scale bar: 200 μm.

Circularity was measured to assess the quality of EBs (figure 6(D)). For the 0.5 mln mL−1 group, most of the EBs were close to a standard spheroid with the circularity centered in 0.9 for the three time points. As to the other two groups, the circularity at day 3 is similar to that of 0.5 mln mL−1group, while the circularity frequency peaks had a significant decrease at day 5 and day 7. In particular, about 20% EBs had a circularity under 0.8 at day 5 and day 7 for the 2.0 mln mL−1group. In general, the circularity decreased with the increase of culture time and initial cell density in the hydrogel (figure 6(E)).

3.6. Comparison with other EB formation methods

Considering this was a novel methodology of EB formation, we systematically compared the commonly used static suspension and hanging drop methods with the 3D bioprinting method for EB formation. As demonstrated by the phase-contrast images (figure 7(A)), EBs generated by static suspension method showed more uncontrollable morphology rather than round spheroid. The distribution of EB diameter clearly demonstrated that 3D bioprinting technology generated EBs with higher uniformity compared with static suspension technology, especially for the larger EB diameter, i.e. 60 ~ 70 μm and 100 ~ 110 μm regions (figure 7(B)). In particular, the EBs with 30 ~ 50 μm diameter presented vastly irregular shape in suspension technology, which was confirmed by the circularity curve (figure 7(C)). On the other hand, EBs generated by 3D bioprinting technology showed higher circularity regardless of the diameter regions, suggesting more regular shape (figure 7(C)). More characteristic like EB forming motivation, size control method, EB diameter range, uniformity, yield, operation complexity were compared among 3D bioprinting technology, static suspension technology and hanging drop technology, as listed in table 1.

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Figure 7. Comparison of static suspension and 3D bioprinting technology for generating EBs. (A) Phase-contrast images showing the morphology of EBs generated by static suspension technology and 3D bioprinting technology. (B) The EB diameter histograms presented the distribution of EB size with a Gauss distribution fitting. (C) The circularity curves contrasted the EB qualities.

Table 1.  Comparison of three EB forming methods.
Hanging-drop Suspension 3D print
Forming mechanism Aggregation by gravity Self-aggregation Proliferation
Size control Time and cell density Time and cell density Mainly time
Diameter range 50 ~ 500 μm 50 ~ 500 μm 30 ~ 200 μm
Uniformity High Low Medium-high
Yield Low High High
Operation Time-consuming for seeding and medium refresh Complex for medium refresh Time-saving and easy for medium refresh


4. Discussion

3D cell culture environment and tissue-like models have drawn great attention because they can be tuned to promote certain levels of cell differentiation and tissue organization, which is difficult in traditional 2D culture systems for their failing to reconstitute the in vivo cellular microenvironment [30, 31]. Various 3D culture systems have been developed to study the cellular behavior affected by spatial and temporal cell–cell and cell–matrix interactions. Among these methods, 3D bioprinting, typically containing jet-, laser- and extrusion-based methods, is a promising technique to manipulate cells/matrix deposition and ultimately generate 3D complex tissues or organs. This technique have been used in printing cells derived from adult, embryonic and even tumor tissues for tissue engineering and drug screening applications. With the capacity to expand unlimitedly in vitro and differentiate into a variety of therapeutic cell types, ESCs have generated great enthusiasm and are being applied in bioprinting studies until recently. As a relatively sensitive cell type, ESCs might suffer greater problems in a printing process compared with other types of cells. Several studies had been conducted to print ESCs, maintaining their viability and pluripotency [1619]. Instead of creating 3D tissue-like constructs, these studies were more likely to generate cellular droplet array with precise control of distribution. Here we described the work of establishing a 3D ESC-laden hydrogel construct using extrusion-based bioprinting technology. The results demonstrated high proliferation rate of pluripotent ESCs in the hydrogel construct, and a versatile technology for generating highly uniform and high throughput EBs.

Cell viability after 3D bioprinting and construct formation was determined when evaluating the limitations of bioprinting ESCs. Cells would be lysed or damaged due to osmotic effects in the solution, heat increase and mechanical stress during printing. In the protocol presented in this work, about 6.86% ± 1.31% cells were dead during the cell/hydrogel solution preparation process before 3D bioprinting (figure 2(B)). We assumed this was caused by cell dissociation process, together with the osmosis and stirring operation of hydrogel materials. In an inkjet printing study, 15% Chinese Hamster Ovary cells were detected dead before printing process [32]. Thermal effects of the ejector reservoir in the inkjet printing process and laser force in laser-based printing would be the cause of cell death, in addition to the impact force when cellular droplets were jetted to a rigid substrate in a very short time. Under a different fabricating strategy, the extrusion-based bioprinter extruded the cell-laden cylinders softly on the substrate and controlled the temperature under 30 °C, without the concerns about the thermal and sharply impacting effects. However, cells would inevitably suffer from shear force when the cell-laden hydrogels were continuously extruded through a limited space in the nozzle. We hypothesized that nozzle size and hydrogel viscosity would influence shear force and hence influence cell viability. The cell viability data of different nozzle sizes, chamber and insulation temperatures supported this hypothesis (figures 3(A) and (B)). In our previous study, more than 90% Hela cells were alive after bioprinting under the parameters of Insu-25 °C/Cham-4 °C and Nozzle-260 μm [23], while the viability of ESCs was only 55.52% ± 2.37% under the same parameter combination. When increasing the insulation and chamber temperature to 30 °C and 10 °C respectively, the viability showed a significant increase to 90%. Taking into the account of cell death before bioprinting, optimized parameters led to only 5% cell death during printing, indicating a broad future applicability of this technique to various cell types ranging from tumor cells to ESCs. Additionally, few dead cells were observed during one-week culture period (figure3(C)). On the other hand, when the culture period was extended to more than 7 days, more and more ESCs suffered from apoptosis and lysis, possibly due to contact inhabitation and insufficient mass transfer to the center of EB with the increasing of EB size. Therefore, 7 days was chose as the experiment time window for this study.

Apart from cell viability, the maintenance of pluripotency is another essential criterion for ESCs regulation and application. The results of immunofluorescence staining and FACS analysis showed a high expression rate (98%) of stem cell pluripotent markers Oct4 and SSEA1 at day 7 (figure 4), indicating that cells remained undifferentiated state during the whole experimental period. Naturally, it can be inferred that the printing process also had little influence on ESC pluripotency.

In the cell-laden hydrogel culture system, both the cell type and matrix material could influence cell growth. Human mesenchymal stem cells remained alive but did not proliferate when encapsulated in alginate [33, 34]. While human ADSCs could proliferated for a short period of time in alginate hydrogel microspheres but showed significantly higher proliferation rate in gelatin/alginate microspheres [35]. As a widely used hydrogel, alginate has the disadvantages of low cell adhesiveness and poor support for cell proliferation [36]. Adding gelatin would improve the cellular adhesive condition and hence favor cell expansion. In this study, the fabricated multilayered constructs offered a 3D microenvironment surrounded by gelatin/alginate materials for ESCs to adhere, self-renew, and cellular spheroid, termed EB, was generated in situ because of cell proliferation. Once EB was formed, the spheroid structure supported expansion of subpopulations with differing proliferation, nutrition and oxygenation status compared with conventional monolayer system. It is reported that the proliferation of mouse ESCs was higher when embedded in fibrin gels versus 2D suspension culture [27]. Similarly, in this study, ESCs in 3D constructs proliferated faster than 2D culture sample when being released from hydrogel to read OD value. This operation was aimed to avoid the influence of interactions between reagent molecular and matrix materials (figure 6 and supplement 3). Additionally, the enlargement of EB diameter, which also reflected ESC proliferation, confirmed this result (figure 6).

Typically stimulated via generation of EBs, ESC differentiation depends on numerous cues throughout the EB environment, including EB size and shape, as well as their uniformities. In general, several characteristics should be concerned for EB formation system, including reproducibility, symmetry, ease of use and scalability [37]. In the traditional EB formation methodology, like suspension and hanging-drop, EBs were created via cell gathering and proliferation. In these methods, it was essential to get a balance between allowing necessary ESC aggregation for EB formation and preventing EB agglomeration for efficient cell growth and differentiation [14]. Static suspension cultures produced a large number of EBs with simple operation, but the size and shape of the resulting EBs were highly uncontrollable and irregular due to the tendency of EBs to agglomerate after initial formation, as shown in figure 7. Hanging-drop method served as a golden tool to generate uniform and reproducible EBs with fully aggregating of cells under gravity and non-agglomeration of EBs in different drops. However, it faced the intrinsic limitation of scalability. The 3D bioprinting method presented in this study addressed some of the problems, producing massively homogeneous EBs with regular shape and controllable shape. In this 3D cell-laden hydrogel system, ESCs were immobilized and restricted to aggregate with each other, and would not agglomerate until they are large enough to connect with each other. When the initial cell density was increased, the average distance between two original EBs was closer and these EBs are more likely to agglomerate with each other while proliferation, which is also one of the concerns when we choose the experiment time period. As a result, the EB uniformity of 2.0 mln mL−1 group was not that good as those of 0.5 mln mL−1 and 1.0 mln mL−1 groups, especially after culturing for one week (figure 6). Without the initial cell aggregating, the size of EBs in our model was mainly determined by the culture time. Also, it would take longer to reach the same scale of EB diameter compared with suspension method, probably due to the physical constrain of the matrix material. For example, it took 5 days and 2 days to get EBs ranging 60 ~ 70 μm for 3D printing and suspension methods, respectively (figure7). Besides, thanks to the interconnected channels design in the 3D construct which allowed mass transfer, EBs could be produced in a large scale by changing the construct volume and cell density. In the six-layer construct with 1.0 million cells per milliliter for example, EBs got a stable yield of about 3000 cm−2, while the EB yield by suspension technology was about 900 cm−2(seeding 0.5 million cells in a 35 mm dish) and no more than 10 EBs [38] could be produced in 1 cm2 area in hanging-drop method, which was also demonstrated by our experiments (supplement 4).

In summary, this study presented the high throughput production of pluripotent, uniform, regular and controllable EBs with the diameter smaller than 150 μm during one week culture. In a gelatin-based laser printing method, EBs with the diameter of about 100 μm were also generated to avoid EB agglomeration in gels [19]. EBs with different size exhibit different gene expression and differentiation fate. Park et al [39] found that 100 μm diameter EBs of mouse ESCs expressed increased ectoderm markers while 500 μm diameter EBs expressed endoderm and mesoderm markers. Furthermore, Messana et al [12] demonstrated that mouse ESCs derived from small EBs (<100 μm) had a greater chondrogenic potential than those from larger EBs. Hwang [10] reported that human endothelial cell differentiation was increased in smaller EBs (150 μm) while cardiogenesis was enhanced in larger EBs (450 μm). However, large EBs might be associated with limited mass transfer and the diffusion of biochemical through EBs is demonstrated to be linked to differentiation of ESCs [40]. While the effect of EB size on differentiation remains to be shown in our model, we hypothesize that EBs with the diameter smaller than 150 μm would mediate specific differentiation trajectory, which will be confirmed in the future work.

Demonstrating the advantages of reproducibility, high throughput, regular shape and controlled size, we believe this is a versatile technology for EB generation. But, this 3D printing system does not serve as an EB formation method solely. The ESC-laden hydrogel 3D construct can be dissolved at a proper time point to harvest massive EBs with desired size for ES cell research. Or, the ESC-laden hydrogel 3D construct can be maintained to perform 3D ESC differentiation studies to explore the regulation of EB size, matrix material and 3D structure on ESC differentiation lineages. Furthermore, this technology hold the potential to serve as a versatile tool for the generation of tissue-like structure and organ/tissue on chip based on controlled ESC differentiation.

5. Conclusion

In this study, we reported successful bioprinting of mouse ESCs with hydrogel into a 3D multilayered construct for the first time. Extrusion-based bioprinting technology was applied. Upon parameter optimization, ESCs demonstrated high viability of 90% after 3D printing and construct formation. Cells continued self-renewal in the construct and exhibited a higher proliferation rate compared with conventional 2D culture. 98% cells expressed the canonical pulripotent markers Oct4 and SSEA1 at day 7, indicating that most of the ESCs remained undifferentiated state after printing and culturing. Large quantities of uniform EBs with regular shape and adjustable size were generated through cell proliferation, while avoiding EBs agglomeration. This work indicated the feasibility of fabricating complex 3D tissue-like model based on pluripotent stem cells for applications in pharmacy, regenerative medicine, stem cell expansion and biology studies.


Bioprinting of human pluripotent stem cells and their directed differentiation into hepatocyte-like cells for the generation of mini-livers in 3D

Alan Faulkner-Jones1,2, Catherine Fyfe3, Dirk-Jan Cornelissen1,2, John Gardner3, Jason King3,4,Aidan Courtney3,4 and Wenmiao Shu1,2

We report the first investigation into the bioprinting of human induced pluripotent stem cells (hiPSCs), their response to a valve-based printing process as well as their post-printing differentiation into hepatocyte-like cells (HLCs). HLCs differentiated from both hiPSCs and human embryonic stem cells (hESCs) sources were bioprinted and examined for the presence of hepatic markers to further validate the compatibility of the valve-based bioprinting process with fragile cell transfer. Examined cells were positive for nuclear factor 4 alpha and were demonstrated to secrete albumin and have morphology that was also found to be similar to that of hepatocytes. Both hESC and hiPSC lines were tested for post-printing viability and pluripotency and were found to have negligible difference in terms of viability and pluripotency between the printed and non-printed cells. hESC-derived HLCs were 3D printed using alginate hydrogel matrix and tested for viability and albumin secretion during the remaining differentiation and were found to be hepatic in nature. 3D printed with 40-layer of HLC-containing alginate structures reached peak albumin secretion at day 21 of the differentiation protocol. This work demonstrates that the valve-based printing process is gentle enough to print human pluripotent stem cells (hPSCs) (both hESCs and hiPSCs) while either maintaining their pluripotency or directing their differentiation into specific lineages. The ability to bioprint hPSCs will pave the way for producing organs or tissues on demand from patient specific cells which could be used for animal-free drug development and personalized medicine.


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New drug development can take 10 to 20 years with an estimated average of about 9 to 12 years [1, 2]. In addition, only around 16% of the drugs that begin preclinical testing are approved for human use [3]. Some of this low success rate can be attributed to the different responses that animals and humans have to the drugs being tested; some drugs have to be withdrawn from market due to toxic effects on human organs such as liver and heart, despite being tested safely on animals. A possible solution to this might be the creation of human pluripotent stem cell (hPSC) -derived micro-tissues which could be used with organ-on-a-chip devices [47]. These micro-tissues are expected to produce the same or similar physiological reaction that the entire organ would but on a much smaller scale. This would result in scalable, faster and potentially more reliable drug testing platform, and hopefully an end to animal testing.

hPSCs are the ideal cells to use for this application due to their ability to self-renew indefinitely, which enables large populations of cells to be created easily in vitro, and their pluripotency which means that they can be differentiated into any required adult cell type [813]. Pluripotent stem cells can be divided into embryonic stem cells (ESCs) and induced pluripotent stem cells (iPSCs). Human ESCs (hESCs) were first isolated from early human blastocysts in 1998 [14]. Any tissue construct created from hESCs for implantation in vivo would require the patient to receive immunosuppressive drugs and ethical issues still restrict some applications due to their source. iPSCs have neither of these drawbacks as they can be created from harvested adult cells from the patient requiring treatment and as such any implanted cells derived from these iPSCs should not be rejected by the patient’s immune system but may require immunosuppressive drugs at a greatly reduced dosage. In 2006 Shinya Yamanaka discovered that iPSCs can be derived from somatic cells by retrovirally transducing them with four transcription factors—Oct3/4, Sox2, Klf4 and C-myc [15, 16]. These cells have the same self-renewal and differentiation capabilities as ESCs but with the added advantage that iPSCs can be used for autologous therapies. These unique characteristics make pluripotent stem cells ideal for use in a number of applications such as clinical tissue engineering, novel drug discovery and testing for the pharmaceutical industry [8,9, 17, 18].

In the field of biofabrication, great advances are being made towards fabricating 3D tissue and organs with very fine spatial control of cell deposition. From the very first paper that was published investigating printing of biological cells (or bioprinting), tissue engineering was identified as a major application for this new technology [19]. If more complex structures such as organs and tissues were to be printed, the bioprinter would need the ability to transfer microscopic patterns of viable cells of multiple cell types into well-defined three-dimensional arrays that closely mimic the tissue structure. There has been much progress in the development and establishment of several different bioprinting techniques for 3D live constructs [2022] including those based on laser pulses, inkjets and other more novel approaches. It is an inescapable fact that cells will be subjected to some level of stress during deposition, regardless of the printing technique being used. For example, cells printed by non-contact methods will be affected when they impact on the substrate at some incident velocity, which would result in extreme deceleration and shear stress [2326]. Shear stress is applied to cells pushed through nozzle orifices and capillary tubes [24, 2742] and the actuation is provided via pressure, heat, or high frequency vibration which can also be damaging to the cells [30, 31, 4346]. If cells are exposed to laser energy the radiation can cause genetic damage [29, 4754] and shear forces are applied during cavitation and jet formation [23, 55]. Ultrasonic actuation for cell transfer would subject the cells to stress in the form of heat and vibration [56, 57]. Therefore, it is important to validate the response of printed cells to any particular bioprinting process in terms of their viability and more importantly their biological functions.

We previously reported the results of the first experiments printing hESCs using a valve-based printing approach including their response to the printing process in the form of post-printed viability and pluripotency validation [37]. However, if hPSCs are to be used for producing human tissues on demand for drug testing, their post-printing differentiation must be reproducibly directed to the required lineages for each tissue. Unfortunately homogenous cellular differentiation of hPSCs into some germ layers has proved difficult [12, 13]. Here, we report the first investigation into the bioprinting of human iPSCs, their response to the valve-based printing process as well as their post-printing differentiation into hepatocyte-like cells (HLCs). HLCs that are in the process of differentiating are bioprinted and examined to further validate the compatibility of the valve-based bioprinting process with fragile cell transfer. Finally, 3D hydrogel structures were designed and printed out with encapsulated hESC-derived HLCs and the viability and hepatic characteristics of the cells were investigated.


A newer version of our previously reported cell printing platform [37] has been developed. Four nanolitre dispensing systems, each comprising a solenoid valve (VHS Nanolitre Dispense Valve, Lee Products Ltd) with 101.6 μm internal diameter nozzles (Minstac Nozzle, Lee Products Ltd), were attached to static pressure reservoirs for the bio-ink solution to be dispensed from via flexible tubing. The nanolitre dispensing system and bio-ink reservoirs were mounted onto the tool head of an enclosed custom built micrometer-resolution 3-axis XY–Z stage (figure 1). This newer cell printing platform improved on the previous version by reducing the overall size and weight of the machine, allowing it to be mounted inside a standard tissue culture hood during experiments requiring a sterile environment. Other enhancements included the two extra nanolitre dispensing systems, taking the total up to four, a more robust electronics and custom firmware was developed which improved the reliability and speed of the machine and two separate pressure channels were included, allowing for differential bio-ink dispensing conditions. Unless otherwise stated standard printing conditions were used: for 2D, printing was carried out using a pulse time of 8 ms at an inlet pressure of 0.6 bar using a nozzle with an internal diameter of 101.6 μm; for 3D, printing was carried out using a pulse time of 400 μs at an inlet pressure of 1.0 bar for sodium alginate solution and a pulse time of 400 μs at an inlet pressure of 0.5 bar for calcium chloride solution both using nozzles with an internal diameter of 101.6 μm.

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Figure 1. (a) Schematic drawing of the cell printer system; (b) detailed schematic of the micro-solenoid valve; (c) schematic of the combinatorial printing process for alginate hydrogel creation; (d) a 3D printed alginate tube structure approximately 13 mm tall printed with 1.5% w/v Sodium Alginate and 600 mM (6%) Calcium Chloride solutions in Millipore water (scale bar 2 mm).



The process of in vivo liver organogenesis occurs in the developing foregut, when newly specified hepatic cells separate from the endodermal sheet and form a dense 3D structure known as a hepatoblast (liver bud) [74, 75]. It is hypothesized that arranging the hESC–HLCs in 3D during the differentiation process may yield more mature hepatocytes than conventional 2D differentiation. The hESC differentiation protocols are more efficient and robust than hiPSC protocols therefore only hESC-derived HLCs were printed in 3D.

In order for this technique to be useful for tissue engineering applications, structures need to be tall enough to allow cells to interact in a three-dimensional environment. The concentration of alginate solution was set to 1.5% w/v to improve the mechanical strength of the hydrogel and allow it to support further layers. Circular structures with a large number of layers were designed and printed out in the wells of a multi-well plate to allow the structures to be cultured post-printing. These resulting structures were photographed for analysis and are shown in figure 7below.

These structures were printed out in a matter of minutes and are strong enough to support their own weight and the weight of further layers (as seen in figure 1(d)). The structures spread slightly, but by slightly altering the volume ratio, concentrations and surface properties this spreading can be reduced.

Approximately one hour post-printing one of the HLC-laden alginate ring structures was examined using a confocal microscope; the 3D image is shown in figure 8(a). Cell viability was calculated to be 55.5% using the Imaris confocal microscope software. Cell viability declined over the first 24 h which resulted in low cell numbers for hepatic marker testing following the 3D differentiation process, but the viability remained stable for the remainder of the differentiation process. At day 23 of the differentiation process, the cells in the remaining structures were harvested and stained for the presence of hepatic markers. As shown in figure 8(b), cells are positive for albumin which demonstrates their hepatic lineage. The normal time required for 2D differentiation of hPSC-HLCs is 17–24 d. However, based on the results of albumin secretion in the medium, we observed the 3D printed cells have taken longer to reach the maximum albumin secretion than the 2D control as shown in figure 8(c). Interestingly, when analyzing the difference between 20 and 40 layer printed tube structures, we noticed close-to proportional increase in albumin secretion to the number of layers as shown in figure 8(d). This indicates that the permeability of the alginate hydrogel allows nutrition and differentiation reagents to enter the structure and support 3D differentiation and maturation processes of the cells, regardless of the height of the printed structure.

Research is currently underway including investigations to improve the 3D viability and adjusting the differentiation protocol that may facilitate higher albumin secretion. For example, the optimization of hydrogel formation as well as enhanced cell density may improve the differentiation process for hPSCs in 3D [21, 76, 77].

4. Conclusions

To the best of our knowledge, this study is the first to demonstrate that hiPS cells can be bioprinted without adversely affecting their biological functions including viability and pluripotency. Importantly, we verified that our valve-based printing process is gentle enough to not affect the pluripotency of both hESCs and hiPSCs. A number of different hPSC lines were directed to differentiate into HLCs. Cells were printed during the differentiation process and showed no differences in hepatocyte marker expression and similar morphology when compared to a non-printed control. We previously reported the results of an investigation into the response of hESCs to the valve-based printing process. Here we build on that study, performing a deeper investigation to compare the response of hiPSCs and hESCs to the printing process using flow cytometry. The effect of nozzle geometry was investigated and the effects of nozzle length on the post-printing viability of cells were recorded; longer nozzles lower the post-printing viability of the cells. We printed hESC-derived HLCs in a 3D alginate matrix and tested for viability and hepatic markers during the remaining differentiation and they were found to be hepatic in nature. The ability to bioprint hPSCs while either maintaining their pluripotency or directing their differentiation into specific cell types will pave the way for producing organs or tissues on demand from patient specific cells which could be used for animal-free drug development and personalized medicine.



Large scale industrialized cell expansion: producing the critical raw material for biofabrication processes

Arun Kumar1 and Binil Starly1,2


Cellular biomanufacturing technologies are a critical link to the successful application of cell and scaffold based regenerative therapies, organs-on-chip devices, disease models and any products with living cells contained in them. How do we achieve production level quantities of the key ingredient—’the living cells‘ for all biofabrication processes, including bioprinting and biopatterning? We review key cell expansion based bioreactor operating principles and how 3D culture will play an important role in achieving production quantities of billions to even trillions of anchorage dependent cells. Furthermore, we highlight some of the challenges in the field of cellular biomanufacturing that must be addressed to achieve desired cellular yields while adhering to the key pillars of good manufacturing practices—safety, purity, stability, potency and identity. Biofabrication technologies are uniquely positioned to provide improved 3D culture surfaces for the industrialized production of living cells.

Biofabrication of tissue constructs by 3D bioprinting of cell-laden microcarriers

Riccardo Levato1,2, Jetze Visser3, Josep A Planell1, Elisabeth Engel1,2,4, Jos Malda3,5 andMiguel A Mateos-Timoneda2,1


Bioprinting allows the fabrication of living constructs with custom-made architectures by spatially controlled deposition of multiple bioinks. This is important for the generation of tissue, such as osteochondral tissue, which displays a zonal composition in the cartilage domain supported by the underlying subchondral bone. Challenges in fabricating functional grafts of clinically relevant size include the incorporation of cues to guide specific cell differentiation and the generation of sufficient cells, which is hard to obtain with conventional cell culture techniques. A novel strategy to address these demands is to combine bioprinting with microcarrier technology. This technology allows for the extensive expansion of cells, while they form multi-cellular aggregates, and their phenotype can be controlled. In this work, living constructs were fabricated via bioprinting of cell-laden microcarriers. Mesenchymal stromal cell (MSC)-laden polylactic acid microcarriers, obtained via static culture or spinner flask expansion, were encapsulated in gelatin methacrylamide-gellan gum bioinks, and the printability of the composite material was studied. This bioprinting approach allowed for the fabrication of constructs with high cell concentration and viability. Microcarrier encapsulation improved the compressive modulus of the hydrogel constructs, facilitated cell adhesion, and supported osteogenic differentiation and bone matrix deposition by MSCs. Bilayered osteochondral models were fabricated using microcarrier-laden bioink for the bone compartment. These findings underscore the potential of this new microcarrier-based biofabrication approach for bone and osteochondral constructs.

Microstereolithography and characterization of poly(propylene fumarate)-based drug-loaded microneedle arrays

Yanfeng Lu1, Satya Nymisha Mantha1, Douglas C Crowder2, Sofia Chinchilla2, Kush N Shah2,3,4,Yang H Yun2, Ryan B Wicker5 and Jae-Won Choi1


Drug-loaded microneedle arrays for transdermal delivery of a chemotherapeutic drug were fabricated using multi-material microstereolithography (μSL). These arrays consisted of twenty-five poly(propylene fumarate) (PPF) microneedles, which were precisely orientated on the same polymeric substrate. To control the viscosity and improve the mechanical properties of the PPF, diethyl fumarate (DEF) was mixed with the polymer. Dacarbazine, which is widely used for skin cancer, was uniformly blended into the PPF/DEF solution prior to crosslinking. Each microneedle has a cylindrical base with a height of 700 μm and a conical tip with a height of 300μm. Compression test results and characterization of the elastic moduli of the PPF/DEF (50:50) and PPF/drug mixtures indicated that the failure force was much larger than the theoretical skin insertion force. The release kinetics showed that dacarbazine can be released at a controlled rate for five weeks. The results demonstrated that the PPF-based drug-loaded microneedles are a potential method to treat skin carcinomas. In addition, μSL is an attractive manufacturing technique for biomedical applications, especially for micron-scale manufacturing.

Controlling shape and position of vascular formation in engineered tissues by arbitrary assembly of endothelial cells

Hiroaki Takehara1,4, Katsuhisa Sakaguchi2, Masatoshi Kuroda3, Megumi Muraoka3, Kazuyoshi Itoga1,Teruo Okano1 and Tatsuya Shimizu1


Cellular self-assembly based on cell-to-cell communication is a well-known tissue organizing process in living bodies. Hence, integrating cellular self-assembly processes into tissue engineering is a promising approach to fabricate well-organized functional tissues. In this research, we investigated the capability of endothelial cells (ECs) to control shape and position of vascular formation using arbitral-assembling techniques in three-dimensional engineered tissues. To quantify the degree of migration of ECs in endothelial network formation, image correlation analysis was conducted. Positive correlation between the original positions of arbitrarily assembled ECs and the positions of formed endothelial networks indicated the potential for controlling shape and position of vascular formations in engineered tissues. To demonstrate the feasibility of controlling vascular formations, engineered tissues with vascular networks in triangle and circle patterns were made. The technique reported here employs cellular self-assembly for tissue engineering and is expected to provide fundamental beneficial methods to supply various functional tissues for drug screening and regenerative medicine.

The influence of printing parameters on cell survival rate and printability in microextrusion-based 3D cell printing technology

Yu Zhao1,2, Yang Li1,2, Shuangshuang Mao1,2, Wei Sun1,2,3,4 and Rui Yao1,2

Three-dimensional (3D) cell printing technology has provided a versatile methodology to fabricate cell-laden tissue-like constructs and in vitro tissue/pathological models for tissue engineering, drug testing and screening applications. However, it still remains a challenge to print bioinks with high viscoelasticity to achieve long-term stable structure and maintain high cell survival rate after printing at the same time. In this study, we systematically investigated the influence of 3D cell printing parameters, i.e. composition and concentration of bioink, holding temperature and holding time, on the printability and cell survival rate in microextrusion-based 3D cell printing technology. Rheological measurements were utilized to characterize the viscoelasticity of gelatin-based bioinks. Results demonstrated that the bioink viscoelasticity was increased when increasing the bioink concentration, increasing holding time and decreasing holding temperature below gelation temperature. The decline of cell survival rate after 3D cell printing process was observed when increasing the viscoelasticity of the gelatin-based bioinks. However, different process parameter combinations would result in the similar rheological characteristics and thus showed similar cell survival rate after 3D bioprinting process. On the other hand, bioink viscoelasticity should also reach a certain point to ensure good printability and shape fidelity. At last, we proposed a protocol for 3D bioprinting of temperature-sensitive gelatin-based hydrogel bioinks with both high cell survival rate and good printability. This research would be useful for biofabrication researchers to adjust the 3D bioprinting process parameters quickly and as a referable template for designing new bioinks.

A new method of fabricating a blend scaffold using an indirect three-dimensional printing technique

Jin Woo Jung1,3, Hyungseok Lee1,3, Jung Min Hong1, Jeong Hun Park1, Jung Hee Shim2, Tae Hyun Choi2and Dong-Woo Cho1

Due to its simplicity and effectiveness, the physical blending of polymers is considered to be a practical strategy for developing a versatile scaffold having desirable mechanical and biochemical properties. In the present work, an indirect three-dimensional (i3D) printing technique was proposed to fabricate a 3D free-form scaffold using a blend of immiscible materials, such as polycaprolactone (PCL) and gelatin. The i3D printing technique includes 3D printing of a mold and a sacrificial molding process. PCL/chloroform and gelatin/water were physically mixed to prepare the blend solution, which was subsequently injected into the cavity of a 3D printed mold. After solvent removal and gelatin cross-linking, the mold was dissolved to obtain a PCL–gelatin (PG) scaffold, with a specific 3D structure. Scanning electron microscopy and Fourier transform infrared spectroscopy analysis indicated that PCL masses and gelatin fibers in the PG scaffold homogenously coexisted without chemical bonding. Compression tests confirmed that gelatin incorporation into the PCL enhanced its mechanical flexibility and softness, to the point of being suitable for soft-tissue engineering, as opposed to pure PCL. Human adipose-derived stem cells, cultured on a PG scaffold, exhibited enhanced in vitro chondrogenic differentiation and tissue formation, compared with those on a PCL scaffold. The i3D printing technique can be used to blend a variety of materials, facilitating 3D scaffold fabrication for specific tissue regeneration. Furthermore, this convenient and versatile technique may lead to wider application of 3D printing in tissue engineering.


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3D-printed organ research enhanced with programmable DNA “smart glue”

Reporter: Irina Robu, PhD
A new breakthrough using DNA to provide the “glue” in a 3D printing material   was created by Dr. Andrew D. Ellington at University of Texas, which can be used to 3D print tissues to repair injuries or even create organs. Since DNA provides the source code for life, the researchers at University of Texas coated plastic microparticles with  40 base pairs of DNA, forming gel-like materials that they could extrude from a 3D printer* to form solid shapes (up to centimeters in size). These were used as scaffolds to host growing cells within the matrix.
The researchers designed DNA sequences which are complementary to one another and which are subsequently fragmented and attached to micro-beads and then anneal to one another and stiffen into a gel-like colloidal structure. The DNA-coated beads could carry and physically orient molecules that do more than simple annealing. The beads can also be coated with an increasing density of a growth factor leading a cell type to grow along the concentration gradient. 
In theory, this could lead to a future in which your outgoing organ is scanned and then its used to  design the large-scale structures of the printing scaffold. The small-scale structure of the tissues will be roughly the same for all livers, and the combination of beads necessary to create it sits ready in the laboratory freezer. Feed the scan and the right beads to the printer, along with samples of liver cells to be deposited in the gel as it is laid down. Print. Wait. Surgery.In addition to the benefits, there are also potential problems since DNA is fragile and the DNA interactions might not last enough for the organ to mature or the slow forming chemical bonds could lock a configuration in place once the DNA had figured it out transiently.As researchers customize XNAs to better offset DNA’s less helpful attributes, the potential will continue to grow.

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Introduction to Tissue Engineering; Nanotechnology applications

Author, Editor and Curator:  Tilda Barliya, PhD


Tissue Engineering is an emerging multidisciplinary field involving biology, medicine, and engineering that is likely to revolutionize the ways we improve the health and quality of life for millions of people by restoring, maintaining, or enhancing tissue and organ function. Tissue engineering emerged as organ transplantation is limited by the number of  available donors and high cost process, leaving thousands of people each year on the transplant waiting lists in the United States alone. Many die before an organ donor becomes available. Dr. Tal Dvir from the Langer’s lab at MIT have summarized this topic in his review (2. http://nextbigfuture.com/2011/01/nanotechnology-strategies-for-tissue.html)

Tissue engineering aims at developing functional substitutes for damaged tissues and organs, a process that involves the use of a combination of cells, engineering and material methods, including suitable biochemical and chemical factors to improve or replace biological functions. Rather than simply introducing cells into a diseased area to repopulate a defect and/or restore function, in tissue engineering the cells are often seeded in or onto biomaterials (scaffolds) before transplantation.

These biomaterial scaffold allows cells to attach and reorganize to form functional tissue by proliferating, synthesizing extracellular matrix, and migrating along the implant path (1,2,3) Figure 1.

Until recently, it was believed that the macroporous features of scaffolds used in tissue engineering mimicked the dimension scale of the extracellular matrix (ECM), and that the matrix itself (natural or artificial) only served as a support for the cells; morphogenesis was controlled passively by defining tissue boundaries. Emphasis was placed on critical engineering and material issues, such as improving mass transfer into the core of the cell constructs and designing biocompatible and biodegradable scaffolds with mechanical properties suitable for engineering various tissues. As the field evolved, attention focused on the biology of the scaffolds and how they affect various cell types.

Tissue engineers had recognized that some of the widely used scaffolds do not fairly recapitulate the cell microenvironment and that the ECM is a dynamic and hierarchically organized nanocomposite that regulates essential cellular functions such as:

  • morphogenesis,
  • differentiation
  • proliferation
  • adhesion
  • migration

Nanotechnological tools for tissue engineering may help design advanced nanocomposite scaffolds that can better mimic the ECM and eventually assemble more complex and larger functional tissues. In order to generate a functional tissue, effective organization of cells in the tissue is required with similar morphology and physiology of the parental tissue.

Morphogenesis in the three-dimensional (3D) scaffold should occur in a similar way to natural organ development. The cells reorganize owing to interaction with the ECM on the basis of:

  • topography,
  • mechanical properties (such as matrix stiffness, elasticity and viscosity)
  • concentration gradients of immobilized growth factors
  • ECM molecules.

Recently, Ott and co-workers (4) reported a study emphasizing the importance of the ECM structure in guiding the seeded cells and promoting morphogenesis. Rat hearts were decellularized by perfusion of detergents to preserve the underlying ECM and then reseeded with cardiac and endothelial cells (4). The cells migrated and self-organized in their natural location in the matrix and by day 8, under physiological load and electrical stimulation, the constructs were able to generate pump function (4). The importance of the ECM was shown for:

  • Heart
  • Lung
  • Arteries
  • Liver
  • Bone
  • Nerve

So why is the Extracellular Martix (ECM) so important?

The ECM is composed of an intricate interweaving of protein fibres such as fibrillar collagens and elastins, ranging from 10 to several hundreds of nanometres. The mesh is covered with nanoscale adhesive proteins such as laminin and fibronectin that provide specific binding sites for cell adhesion (interacting with integrins, cadherins and so forth) and have been shown to regulate important cell behaviours such as growth, shape, migration and differentiation. Polysaccharides such as hyaluronic acid and heparan sulphate fill the interstitial space between the fibres and act as a compression buffer against the stress placed on the ECM or serve as a growth factor depot (Figure 2).

Scaffold design considerations

The ECMs of various tissues in the body differ in the composition and spatial organization of the collagens, elastins, proteoglycans and adhesion molecules, to maintain specific tissue morphologies and organ specific shape and function, and to supply specific instructive cues. Therefore, the design considerations for scaffolds should vary according to the desired engineered tissue. For example, the biochemical, electrical and mechanical functions of the heart are uniquely dependent on their biological nanostructures. The heart’s 3D ECM network is composed of an intricate, micro- and nanoscale interweaving pattern of fibrillar collagen and elastin bundles that form a dense, elastic network with proteoglycans and with adhesive and non-adhesive molecules. In this defined mesh, the cardiomyocytes are forced to couple mechanically to each other, to form elongated and aligned cell bundles that interact with each other or with neighbouring capillaries and nerves.

Post-isolation cells lose their ultrastructural elongated morphology and their interaction with their surroundings, and they adopt a random distribution on the flat surface of the scaffold, which compromises many of their physiological properties. Therefore, the structure and support of the ECM is crucial. See Figure 2.

Limitations of the ECM:

  • Weak mechanical properties
  • Lack of electrical conductivity
  • Absence of adhesive and micoenvironment- defining moieties
  • Inability of cells to self-assemble to 3D tissue structure.

The rational behind incorporating nanostructures is to compensate for other scaffold limitations (Table 2) Ref.2

The Heart for example requires more than alignment and mechanical support (Boyang Zhang, Ref 5)

  1. cell responses to micro- and nanopatterned topographical cues
  2. cell responses to patterned biochemical cues
  3. controlled 3D scaffolds
  4. patterned tissue vascularization
  5. electromechanical regulation (conductivity). of tissue assembly and function

Nanostructures can be used to record the electronic signals that are transmitted through cells such as neurons and cardiomyocytes. One way to record these signals is by lithographically defining nanostructures as field-effect transistors, which are sensitive to local electric field changes. In particular, silicon nanowire transistors are useful for measuring extracellular signals because they exhibit particularly exquisite field-effect sensitivity compared with conventional, planar devices; they are just tens of nanometres in diameter and can therefore interface with cells and tissue at a subcellular level; and they show nanotopographic features that encourage tight interfaces with biological systems.



This introduction reviewed some of the aspects required for tissue engineering  with the affiliation to nanotechnology. In the next post, we will dive deeper into a specific tissue organ, the bioengineering aspect and how nanotechnology strategies may improve the design and outcome.



1. http://www.nanotech-now.com/news.cgi?story_id=35168

2.  Dvir T.,  Timko BR., Kohane DS., and Langer R. Nanotechnological strategies for engineering complex tissues. Nature Nanotechnology 2010; 12():. http://nextbigfuture.com/2011/01/nanotechnology-strategies-for-tissue.html

3. http://www.nature.com/nnano/journal/v6/n1/abs/nnano.2010.246.html

4.  Ott, H. C. et al. Perfusion-decellularized matrix: Using nature’s platform to engineer a bioartificial heart. Nature Med. 14, 213–221 (2008).

5. Boyang Zhang, Yun Xiao, Anne Hsieh, Nimalan Thavandiran and Milica Radisic. Micro- and nanotechnology in cardiovascular tissue engineering. Nanotechnology 2011; 22(49): 494003

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