Bioprinting living structures

Vladimir Mironov,a Glenn Prestwichb and Gabor Forgacscde

Received 7th December 2006, Accepted 26th March 2007

First published as an Advance Article on the web 18th April 2007

DOI: 10.1039/b617903g

Present efforts in tissue engineering are aimed at building living structures by employing the self-

organizing properties of cells and tissues and automated technologies. One such technology is

bioprinting that utilizes three-dimensional delivery devices for the rapid and accurate placement

of biological materials into biocompatible environments, where post-printing self-assembly takes

place. This Application article summarizes the scientific basis of this approach and some of the

recent developments.

1. Introduction

On any given day in the United States about 60 000 people are

waiting for kidney transplants, 3 000 for heart transplants and

17 000 for liver transplants. A number of those on the waiting

list die before an appropriate donor can be found. These

numbers illustrate that the chronic lack of replacement organs

presents a burning problem and if no solutions are soon found,

with increasing life expectancy, the situation will become

critical in the near future. None of the approaches so far

employed seem to provide an answer: the number of cadavers

whose body parts are available for transplantations is far

below demand, most mechanical devices (i.e. artificial heart)

fail, xenotransplantation, the use of animal organs, raises

serious ethical issues. Recent developments give hope that

tissue engineering may offer at least a partial remedy.

Tissue engineering is an interdisciplinary field that applies

and combines the principles of engineering and life sciences

toward the development of biological substitutes that restore,

maintain, or improve tissue function or a whole organ.1–7 The

basic idea underlying classical tissue engineering is to seed

living cells into a biocompatible and eventually biodegradable

environment (the scaffold), and then culture this construct in a

bioreactor so that the initial cell population can expand into a

tissue.8 With an appropriate scaffold that mimics the

biological extracellular matrix (ECM), the developing tissue

will adopt both the form and the function of the desired organ.

aDepartment of Cell Biology and Anatomy, Medical University of SouthCarolina, Charleston, SC 29425, USAbDepartment of Medicinal Chemistry, The University of Utah, Salt LakeCity, UT 84108, USAcDepartment of Physics, University of Missouri, Columbia, MO 65211,USAdDepartment of Biology, University of Missouri, Columbia, MO 65211,USAeDepartment of Bioengineering, University of Missouri, Columbia,MO 65211, USA

Vladimir Mironov

Vladimir Mironov, MD, PhDis Associate Professor at TheDepartment of Cell Biologyand Anatomy and Director ofBioprinting Research Center atThe Medical University ofSouth Carolina. His researchinterests include cardiovasculardevelopmental biology and tis-sue engineering. Dr Mironov isVice President of the WorldAcademy for Bioprinting.

Glenn D. Prestwich

Glenn D. Prestwich isPresidential Professor ofMedicinal Chemistry at TheUniversity of Utah. He receivedhis PhD from StanfordUniveristy. His main researchinterest is in designing andengineering synthetic extra-cellular matrices for tissueengineering applications. He isco-founder and scientific advisorfor Echelon Biosciences, SentrxSurgical, Carbylan BioSurgery,Sentrx Animal Care, andGlycosan BioSystems. Gabor Forgacs

Gabor Forgacs is GeorgeH. Vineyard Professor ofBiophysics at the University ofMissouri. He received his PhDin theoretical physics from theRoland Eotvos University,Budapest, Hungary. His pre-sent research interest is in thephysical mechanisms that act inearly embryonic development.He is applying these mechan-isms to building living struc-tures of prescribed shape bybioprinting.

APPLICATION www.rsc.org/materials | Journal of Materials Chemistry

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The organ would then be implanted into the recipient.2 This

approach has lead to some spectacular results.9 It has recently

been reported that several patients received tissue engineered

bladders, constructed from their own cells (autologous tissue

engineering), which resulted in regained bowel function.10

Despite these successes, the classical tissue engineering appro-

ach faces serious hurdles. Selection of the ideal biomaterial

scaffold for a given cell type has been accomplished to date by

trial and error. At least three key parameters must be adjusted

in order to optimize the biomaterial. First, a cell-specific

composition of the scaffold, including ECM proteins and

glycosaminoglycans, and necessary growth factors must be

determined. Second, the degradation process of the biomate-

rials on one hand must be sufficiently long to support cells

during proliferation and differentiation, and on the other hand

short enough to be permissive for tissue growth. Third, the

compliance of the biomaterial must be matched to the cell

type; soft tissues will require biomaterials with less rigidity

than hard tissues. Within these three parameters, structural

integrity must be maintained during organogenesis. More

importantly, when several cell types are present, as is the case

for virtually all organs, a new optimization process may be

required. Pre-shaping the scaffold may present further

difficulties. This Application article reviews some novel

approaches in the field, in particular self-assembly-based tissue

engineering and its adaptation to the technology of three-

dimensional bioprinting.

2. Self-assembly-based computer aided tissueengineering

Self-assembly is a fundamental process that drives structural

organization in both inanimate and living systems.11 It is in the

course of self-assembly of cells and tissues in early develop-

ment that the organism and its parts eventually acquire their

final shape. Even though developmental patterning through

self-assembly is under strict genetic control it is clear that

ultimately it is physical mechanisms that bring about the

complex structures. Self-assembly-based tissue engineering

aims at utilizing the inherent organizational capacity of cells

into tissues and eventually organs by mimicking natural

morphogenesis. For example, endothelial cells are genetically

predestined to form blood vessels (depending on the diameter,

possibly in combination with other cell types). Thus, these cells

will form tubular structures on their own, provided the right

external conditions are assured. In principle, there is no need

to pre-shape any scaffold to arrive at such a structure. The

primary challenge will be to identify and in vitro provide the

correct external conditions.

It has recently been suggested that the process of building

three-dimensional (3D) biological structures with prescribed

geometry could significantly be accelerated by the technology

of bioprinting: the automated, computer aided, layer-by-layer

deposition of cells and cell aggregates.12 Commercially

available inkjet printers have been successfully re-designed or

new ones built to specifically deliver biological material into

scaffolds fabricated according to a computer-aided design

template.13,14 Pressure-operated mechanical extruders have

been developed to handle biomolecules or live cells.15,16 Both

technologies have their advantages and deficiencies. In the

inkjet technology, cells are exposed to harsh mechanical

conditions and it is questionable that high enough cell densities

can be achieved to arrive at functional tissues. On the other

hand inkjet printers are cheap, fast and versatile. These

printers can be outfitted with numerous nozzles, thus parallel

processing is possible. Mechanical extruders are more gentle

for cells and are capable of delivering multicellular aggregates.

Such devices are more appropriate for building structures that

require high cell densities. However, extruders are more

expensive, slower than inkjet printers and parallel processing

is not simple. It is likely that both technologies will be used in

the near term, and that demand for automated delivery of cells

will lead to a hybrid approach in the future.

We will separate the tissue engineering technologies required

for bioprinting into four components. First, a bio-ink is

needed. In our work, the bio-ink consists of small tissue

fragments in the shape of spherical multicellular aggregates of

variable size and composition as required for a particular

application. Second, the bio-paper component is needed. We

describe a bio-compatible, in situ-crosslinkable mimic of the

extracellular matrix that has been developed for injectable

tissue engineering applications. Third, the bio-printer is

needed. This is a computer-controlled delivery device for 3D

formation of the cellularized construct. Finally, a bioreactor

will be necessary to mature the proto-organ into an

implantable, functional entity. The 3D structures are arrived

at via a three-phase process (Fig. 1): (i) pre-processing, or bio-

ink preparation, (ii) processing, i.e. the actual automated

delivery/printing of the bio-ink particles into the bio-paper by

the bio-printer, and (iii) post-processing, i.e., the maturation/

incubation of the printed construct in the bioreactor.

Final structure formation takes place during post-processing

via the fusion of the bio-ink particles (hence ‘‘ink’’), akin to the

coalescence of liquid droplets (Fig. 1). The ability of the

multicellular aggregates to fuse (the basis of this technology) is

the consequence of tissue liquidity, the apparent similarity

between liquids and tissues composed of motile and adhesive

cells.17,18 Tissue fragments of such cells round up into spheres

to minimize their interfacial area.19–21 Contiguously placed

spheres fuse with the same kinetics as liquid drops.22 Two

randomly intermixed distinct cell populations sort, with the

same time evolution and final configuration as phase separat-

ing immiscible liquids.23 All these phenomena can be inter-

preted in terms of tissue interfacial tensions and viscosities,

which have been measured for a number of cell types and their

values found to be consistent with the mutual sorting behavior

of the corresponding tissues.20,21,24–28 The molecular basis of

tissue liquidity has been established by the differential

adhesion hypothesis (DAH),17 stipulating that it is the distinct

cell adhesion apparatus characterizing cohesive tissues that

gives rise to their surface tensions.18 The predictions of DAH

and tissue liquidity for morphogenetic structure formation

have been experimentally verified.29–31

2.1 Preprocessing: the bio-ink

The bio-ink particles are cellular spheroids (Fig. 1) of desired

composition and size that can be prepared in a number of

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ways. In the hanging drop method cells are placed in a drop of

tissue culture medium in a Petri dish, which is then inverted.

Cells descend to the bottom of the drop due to gravity and form

a spheroid. The size of the spheroid is determined by the number

of cells in the drop. In the re-aggregation method cells from a

confluent dish are collected, transferred to a round bottom glass

tube and centrifuged. The firm thin pellet that forms is removed

from the tube and cut into fragments that are incubated on a

gyratory shaker until they round up. Here the size of the

resulting spheroids depends on the dimensions of the fragments.

Multicellular spheroids can be prepared from a single cell

type (single colored bio-ink) or from a mixture of several cell

types (multicolor bio-ink). The bio-ink particles are subse-

quently packaged into bio-cartridges (micropipettes of appro-

priate diameter), which are stored in a culture medium in

the incubator until use. Standardization of the bio-printing

technology requires uniform bio-ink droplets, which can rapidly

be prepared in large numbers. These conditions can be easily

met. Using for example hanging drops, these are prepared in

multi-well culture dishes by dispensing the same amount of

liquid with a similar number of cells in each individual well.32,33

2.2 Processing: the bio-paper and the bio-printer

A wide variety of biointeractive hydrogels have been developed

for tissue engineering,34,35 tissue repair, and release of drugs

and growth factors,4,36 but not all would be suitable for

preparing a printable bio-paper. The Center for Therapeutic

Biomaterials (CTB) at the University of Utah has been

experimenting with hydrogels based on the extracellular matrix

(ECM), a complex, heterogeneous interacting set of proteins

and glycosaminoglycans (GAGs). In the physiological ECM,

covalent interactions connect chondroitin sulfate (CS),

heparan sulfate (HS) and other sulfated GAGs to core

proteins forming proteoglycans (PGs). Non-covalent interac-

tions include the binding of PGs to hyaluronan (HA),

electrostatic associations, hydration of the polysaccharide

chains and collagen triple helix formation.

As a result of this activity simple and effective biocompa-

tible, in situ-cross-linkable hydrogels were developed that

mimic the natural ECM as the bio-paper into which cells or

cell aggregates can be printed.37 These synthetic ECMs

(sECM) recapitulate the minimal composition required to

allow cell attachment and growth as well as the appropriate

liquidity to permit the fusion of cell aggregates.38 The bio-

paper sECM technology is highly versatile; composition, gel

stiffness, and cross-linking rate can all be independently

optimized. This versatility is crucial in the bio-printing

paradigm, as the cell–gel interfacial parameters determine

fusion according to computational models.15,39 Moreover, the

ability to provide spatiotemporal control of growth factor

availability for TGFb, KGF, VEGF, bFGF, Ang-1, PDGF,

HGF, EGF (standard acronyms used in the life sciences), and

others, enhances the potential of the sECM hydrogels in

facilitating cellular self-organization in the printed structures.

In particular, controlling the release of VEGF, an angiogenic

growth factor40–43 provides the possibility for vascularization

of the printed constructs, a crucial and still not resolved

problem in tissue engineering.

Fig. 2 shows a diagram of a minimalist sECM that was

designed to support cell attachment, growth, and proliferation

in 3D. Gelatin (Gtn) and HA were chemically modified to give

the corresponding thiolated dithiopropionylhydrazide (DTPH)

derivatives, Gtn-DTPH and HA-DTPH. Co-crosslinking Gtn-

DTPH with HA-DTPH (Fig. 2) afforded materials to which

cells readily attached and spread.44,45 This was achieved with

injectable materials using poly(ethylene glycol) diacrylates

(PEGDA) crosslinking.44 In this case, gelatin substitutes for an

RGD peptide46 or a mixture of three recombinant domains of

human fibronectin,47 which both offer sECMs that promote

cell growth in vitro and tissue formation in vivo. An important

feature of these materials in contrast to using collagen gels is

called ‘‘nanostenting’’.48 When fibroblasts in a collagen gel are

activated with platelet-derived growth factor (PDGF), the gel

undergoes contraction to less than 30% of the original volume.

In contrast, when fibroblasts are encapsulated in a crosslinked

sECM containing unmodified collagen, no contraction occurs

following cell activation. This nanostenting feature is crucially

important in printing tissue constructs. These sECM bio-

materials have already been successfully used for delivery of

autologous bone marrow cells for repair of an osteochondral

defect in a rabbit model49 confirming that the sECMs have the

desired biocompatibility, composition, biodegradation, and

compliance required for effective tissue repair. Biochemical

and biophysical optimization of these sECM for a given cell

type or mixtures of several cell types can be achieved by

Fig. 1 Sequence of events in the bioprinting process. A: Schematic.

Spherical bio-ink particles, prepared in the pre-processing phase, are

packaged into bio-cartridges (not shown) and in the processing phase

delivered by the bio-printer into the bio-paper (sheets) layer by layer.

Maturation of the printed structure into the desired form takes place in

the post-processing phase through the fusion of the bio-ink particles.

B: A ring of 12 bio-ink particles, composed of Chinese Hamster Ovary

cells, printed into collagen type 1 (1 mg ml21) hydrogel (left) fuses into

a continuous toroidal structure in about 120 h (right).

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adjusting the ratio of thiol-modified Gtn and HA (Fig. 3), as

well as adding additional proteins or crosslinkable heparin to

mimic heparan sulfate proteoglycans.43

2.3 Post-processing: structure formation through the fusion of

bio-ink particles

Once the bio-ink particles have been delivered into the bio-

paper, the printed construct is transferred into a bioreactor,

which could be a simple incubator or a specifically designed

culture environment that enables the control of environmental

variables that affect biological processes.50 (For example, the

maturation of tissue engineered blood vessels needs to take

place under pulsatile flow in order to assure that the final

product possesses appropriate biophysical characteristics.) The

final structure is formed through the fusion of the bio-ink

particles.

Fig. 4 shows some outcomes of the described bio-printing

process. The ring-like structure shown in Fig. 4 shows the

sensitivity of the fusion process to the properties of the bio-

paper. The figure also shows the result of computer simula-

tions based on a simple tissue-liquidity-based model, described

below. The striking agreement between experiment and model

strongly suggests that tissue liquidity indeed can be the

morphogenetic mechanism underlying post-printing structure

formation. According to Fig. 4, our approach to tissue

engineering of blood vessels takes place through the following

Fig. 2 Top: Schematic of a two-component sECM. Bottom: Chemical structure of crosslinked sECM formed from HA-DTPH, gelatin-DTPH,

and PEGDA.

Fig. 3 Cell motility in Gtn-HA sECMs. The figure shows the

invasion of the sECM by Chinese Hamster Ovary cells starting from

a compact bio-ink particle. For this particular cell type maximum

spreading is achieved for HA concentration around 30%.

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process. Rings of aggregates shown in Fig. 4C, with the

appropriate cellular composition are printed layer-by-layer on

top of each other (see also Fig. 1A). Upon fusion these provide

lumen-containing vessels, which when attached to each other

result in branching structures. The temporary lumen-filling

hydrogel can be removed by special enzymes (e.g. hyaluroni-

dase in the case of hyaluronan hydrogel).

The outlined process to build 3D living structures of specific

shape has the following advantages. First, it is based on solid

scaffold-free tissue engineering. Second, it utilizes the capacity

of robotically placed self-assembling cell aggregates or tissue

spheroids to fuse into 3D tissue constructs. Third, after the

completion of printing, complex tissue built in this way will

have an incorporated (‘‘built in’’) 3D shape; in the particular

case of blood vessels, a branched vascular tree. However, the

biomechanical properties of such vascular constructs would be

inferior to true blood vessels, even if they are supported by

printing parenchymatous tissue to surround them. Thus there

is need for the development of technologies to improve the

material properties of printed solid scaffold-free 3D vascular

tissue constructs. Theoretically, there are at least three main

approaches: genetic, chemical and physical stimulation that

could collectively be employed in ‘‘accelerated vascular tissue

maturation’’. The genetic approach may involve the temporal

transfection of cells with lysyl oxidase or TGFb. It has been

shown in culture that smooth muscle cells transfected with

lysyl oxidase (LOX) or TGFb generate biomechanically

stronger tissue constructs.51–53 It was also shown that the

inhibition of versican synthesis using antisense DNA acceler-

ated the formation of elastic fibers in tissue engineered

constructs.54 Our data demonstrate55 that embryonic cushion

tissue explants transfected with the periostin gene have

improved biomechanical properties, indicating that genetic

manipulations have a direct effect on tissue structural integrity.

Finally, the addition of collagen producing cells such as

fibroblasts into smooth muscle aggregates could potentially

accelerate vascular tissue maturation. In the case of physical

stimulation, vascular tissue maturation inducing factors are

perfused with appropriate perfusion pressure and associated

shear stress. The feasibility of such biomechanical conditioning

for vascular tissue engineering has been previously demon-

strated.56,57 To this end we have developed a perfusion

bioreactor which is adaptable to the perfusion of a printed

3D organoid with one artery and one vein.58 Chemical

approaches towards accelerated tissue maturation include

the use of collagen synthesis, deposition, assembly and

cross-linking inducing factors (e.g. vitamin C, LOX, TGFb)

or non-enzymatic glycation with ribose.59 Thus, accelerated

tissue maturation is a challenging but feasible technological

concept.

Fig. 4 Top: Toroidal structure formation in the experiments (left 3 panels) and in the simulations (right 3 panels). Upper and lower rows

correspond respectively to the initial and final configurations. The final configurations in the experiments and simulations are reached in 144 h and

50 000 Monte Carol steps (MCS), respectively. In panels (A,B) the imbedding gel is agarose, whereas in panels (C,D) and (E,F) it is collagen with

respective concentrations 1.0 mg ml21 and 1.7 mg ml21. The nuclei of the cells have been fluorescently labeled. Since individual cells produce a

weak fluorescent signal, the strong dispersion of cells into the matrix in panel F (similarly to panel I) is not fully visible. The model simulations in

panels (G,H), (I,J) and (K,L) correspond to representative runs performed with ccg/ET = 10, 0.9 and 0.25, respectively. Bottom: Sheet formation

depends on the initial configuration and the tissue–matrix interfacial tension. Two initial states, made of model cell aggregates, 925 cells each,

packed in a hexagonal (a) and square lattice (b), after 250 000 MCS evolve into configurations shown in panels c (ccg/ET = 0.8) and d (ccg/ET = 1.4),

respectively. For identical parameters, fusion from the hexagonal initial configuration is considerably faster. Similar structures of 25 aggregates of

CHO cells (500 micron in diameter) were embedded in 1.0 mg ml21 collagen type I (e and f). Compact sheets after 144 hours of incubation are

shown in panels g and h.

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The printed vascular tree will consist of a series of

harmonically branched segments with progressively reduced

lumenal diameter, wall thickness, number of smooth muscle

cell layers and extracellular matrix structural components

(such as elastin and collagen). The vascular segments will be

embedded into the surrounding printed parenchymatous

tissue. In the case of the large diameter blood vessels

connecting the engineered organ to the body, these vessels

must pass the suture test in order to be FDA approved,

indicating that they must be mechanically sound. These factors

imply that the demand for accelerated tissue maturation and

the associated enhancement of the material properties of the

vascular wall is vital.

2.4 Modeling post-printing self-assembly and structure

formation

To investigate if in the novel tissue engineering approach

post-printing pattern evolution indeed takes place through a

liquid-like mechanism a simple three-dimensional model was

constructed, in which the sites of a cubic lattice are occupied

either by point-like cells or gel volume elements. The total

interaction energy, E, of such a system can be written as

E~X

vr, r0w

J sr, sr0ð Þ, where r and r9 label lattice sites, and ,r, r9.

signifies summation over neighboring sites, each pair counted

once. First, second and third nearest neighbors are included, and it

is assumed that a cell interacts with the same strength with all the

26 cells it comes into contact with in 3D. To specify occupancy, a

variable s is assigned to each lattice site with values 0 for a ‘‘gel

particle’’ and 1 for a cell. The interaction energy of two neighbors,

J(sr,sr9) then may take either of the values J(0,0) = 2egg, J(1,1) =

2ecc or J(0,1) = J(1,0) = 2ecg. Here the positive quantities ecc, egg

and ecg are characteristic material parameters that account for

contact interaction strengths for cell–cell, gel–gel and cell–gel pairs,

respectively. More specifically, these are mechanical works needed

to disrupt the corresponding bonds. (Note that ecc and egg are

works of cohesion, whereas ecg is work of adhesion per bond.60)

The strength of cell–cell interaction depends on the cells’ adhesion

apparatus, whereas the cell–gel and gel–gel interactions depend on

the specific chemistry of the gel and are tunable (for example by

the relative concentration of HA and Gtn in the sECM as shown

in Fig. 3 or by the concentration of collagen as shown in Fig. 4).

The total energy expression, E, can conveniently be rewritten in

terms of the interactions strength by separating interfacial and

bulk terms: E = ccgBcg + const. Here Bcg is the total number of

cell–gel bonds, and ccg = (ecc + egg)/2 2 ecg is proportional to the

cell–gel interfacial tension,60 a quantity characterizing liquids in

contact. (In fact the described model is a lattice-gas model of a

binary liquid.61) The remaining terms in E do not change as the

cellular pattern evolves. This model is inspired by earlier efforts

aiming at computer simulations of cell sorting.62,63

The evolution of the system is followed using Monte Carlo

simulations.64 The program identifies the cells on the

aggregate–gel interface, picks one of them randomly, and

exchanges it with an adjacent gel particle chosen by chance.

The corresponding change in adhesive energy, DE, is

calculated and the new configuration accepted with a

probability P = 1 if DE ¡ 0 or P = exp(2bDE) if DE . 0.

b = 1/ET, is the inverse of the average biological fluctuation

energy ET, analogous to the thermal fluctuation energy,23 kBT

(kB = Boltzmann’s constant, T = absolute temperature). In

statistical mechanics this energy characterizes thermal agita-

tion in a system of atoms or molecules. In the case of cells, it is

a measure of the spontaneous, cytoskeleton driven motion of

cells, able to break adhesive bonds between neighbors via

membrane ruffling,65 or more generally, via membrane

protrusive activity (e.g., filopodial extensions). By definition,

a Monte Carlo step (MCS) or ‘‘unit of time’’, is completed

when each cell in contact with the gel has been given the

chance to move once. During each MCS the interfacial sites

are selected in random order. The gel boundary is treated as a

fixed physical limit of the system, and cells are constrained to

move within the gel. As Fig. 4 shows simulations based on this

simple model surprisingly well describe pattern evolution in the

post-printed cellular system, supporting the assumption that it

proceeds in analogy with liquids. It has to be noted that the

final configurations in Fig. 4 do not correspond to the lowest

energy state of the binary liquid model. As has been discussed

earlier,15 the shown arrangements represent long-lived states,

which eventually collapse into the true lowest energy state of

the model, that being a sphere with minimal interfacial area.

Analogously, the preferred experimental configurations in

Fig. 4 (the fused ring and sheet) are also long-lived cellular

configurations,39 which however can easily be preserved by

the elimination of the embedding hydrogel (e.g. collagen or

hyaluronan by the enzymes collagenase and hyaluronidase,

respectively).

3. Summary and outlook

We have outlined a four-part strategy for organ printing based

on bio-ink, bio-paper, bio-printer, and bioreactor technolo-

gies. Already, important advances have been made in each

technology that provide optimism for the challenges ahead.

While printed organs may not be available to patients in this

decade, the rapid progress in research may result in functional

kidneys, livers, vascular, cardiac, and orthopedic materials in

the decade that follows.

Acknowledgements

We thank the NSF-FIBR program and the Centers of

Excellence Program of the State of Utah for financial support.

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2060 | J. Mater. Chem., 2007, 17, 2054–2060 This journal is � The Royal Society of Chemistry 2007

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