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 TECHNICAL ESSAY #2 Select a topic that is of interest to you, as well as pertinent to the field of Histology, and write a 10-page technical essay. You should use a minimum of at least five references in the preparation of your review essay. LOOk at the attached file for more instruction on how to write the essay 

Komalpreet Kaur

Dr. Flosi

Histo 4310 Tech Essay #2


Topics for background and discussion of the article entitle “ Tissue- Engineered Lungs for in Vivo Implantation”


Background about histology

· The need of Tissue Engineering (Tissue Engineering: current strategies and Future directions)

· Lack of organ donors

· Assist in lung tissue regeneration

· What my paper discusses (Tissue- Engineered Lungs for in Vivo Implantation)

· Lung implants in rats, shows promise but there are a few set backs

· The result of the scaffold that lung stem cells may adhere too

· What is a scaffold and why is it important?

· Important for cell adhesion, specifically stem cells

· Leaves a structure of that still has the extracellular matrix

· Gives way for branching of the lungs, critical for successful lung transplantation.

· Recipient of tissue engineered tracheal replacement (Engineering tissues for children: building grafts that grow)

· Discuss why it worked

· Also discuss how this paper can be used to discuss improvements for the primary source.

· Issues with tissue engineered: biocompatibility

· There could be negative side effects, such as chemical breakdowns that could lead to degrading interactions

· Immunological effects (biocompatibility of implants: lymphocyte/ macrophage interactions)


This article is about using in vitro techniques to regenerate lung tissue from adult rat. Because the only way to repair lung tissues is to do a lung transplants in the human body, which cost money, time and the need for a donor. This new methods has been an advance in technology and science to be able to produce a new lung that is functional (at least in experimental procedure) so far for scientist to expand it to experimentation on human lungs tissue to see what the results would be like. The experiment they did in this article shows, the removal of cellular components and leaving only the extra cellular matrix, it was able to reform its structure and airways and was involved in gas exchanged when put back into the adult rat.


Anderson, J., McNally, A. 2011. Biocompatibility of implants: lymphocyte/ macrophage interactions. Semin Immunopathol. 33:221-223

Elliott MJ, De Coppi P, Speggiorin S., et al. 2012. Stem-cell-based, tissue engineered tracheal replacement in a child: a 2-year follow-up study. Lancet 380(9846): 15-21.

Olson, J., Atala, A., Yoo, J. 2011. Tissue Engineering: Current strategies and Future Directions. Chonnam Medical Journal 47(11): 1-13

Petersen, T., Calle, E., Zhao, L., Lee, E., et al. 2010. Tissue- Engineered Lungs For in Vivo Implantation. Science. 329: 538- 541.

DOI: 10.1126/science.1189345
, 538 (2010);329 Science

et al.Thomas H. Petersen
Tissue-Engineered Lungs for in Vivo Implantation

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Institutes of Health Research (MOP-77639) (A.E.).
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Promotion Foundation, and P.J.C. from the John and
Fannie Hertz Foundation.

Supporting Online Material
Materials and Methods
SOM Text
Figs. S1 to S23
Tables S1 to S6

16 February 2010; accepted 4 June 2010

Tissue-Engineered Lungs for
in Vivo Implantation
Thomas H. Petersen,1,2 Elizabeth A. Calle,1 Liping Zhao,3 Eun Jung Lee,3 Liqiong Gui,3

MichaSam B. Raredon,1 Kseniya Gavrilov,4 Tai Yi,5 Zhen W. Zhuang,6

Christopher Breuer,5 Erica Herzog,6 Laura E. Niklason1,3*

Because adult lung tissue has limited regeneration capacity, lung transplantation is the primary therapy
for severely damaged lungs. To explore whether lung tissue can be regenerated in vitro, we treated lungs
from adult rats using a procedure that removes cellular components but leaves behind a scaffold of
extracellular matrix that retains the hierarchical branching structures of airways and vasculature. We
then used a bioreactor to culture pulmonary epithelium and vascular endothelium on the acellular lung
matrix. The seeded epithelium displayed remarkable hierarchical organization within the matrix, and
the seeded endothelial cells efficiently repopulated the vascular compartment. In vitro, the mechanical
characteristics of the engineered lungs were similar to those of native lung tissue, and when implanted
into rats in vivo for short time intervals (45 to 120 minutes) the engineered lungs participated in gas
exchange. Although representing only an initial step toward the ultimate goal of generating fully
functional lungs in vitro, these results suggest that repopulation of lung matrix is a viable strategy
for lung regeneration.

ung diseases account for some 400,000
deaths annually in the United States (1).
Human lungs do not generally repair or

regenerate beyond the microscopic, cellular level.
Currently, the only way to replace lung tissue is
to perform lung transplantation, which is an ex-
pensive procedure that achieves only 10 to 20%
survival at 10 years and one that is hampered by a
severe shortage of donor organs (2). Recently,
techniques have been developed to quantitatively
decellularize complex organs such as heart, liver,
and kidney (3–5). Acellular matrices can provide
attractive scaffolds for repopulation with lung-
specific cells for lung engineering because the
extracellular matrix template should contain ap-
propriate three-dimensional (3D) architecture and

regional-specific cues for cellular adhesion. To
be functional in vivo, an engineered lung should
(i) contain lung-specific cells, (ii) display the

branching geometry of the airways and contain
a perfusing microvasculature, (iii) provide bar-
rier function to separate blood from air, and (iv)
have mechanical properties that allow ventila-
tion at physiological pressures.

Here, we describe our progress toward the
construction of a functional tissue-engineered
lung, using rat as a model system. Our approach
is summarized in Fig. 1. We first decellularized
native lung tissue in order to remove all immu-
nogenic cellular constituents (Fig. 1, A and B).
We found that after careful decellularization, the
tissue retained its alveolar micro-architecture, its
ability to function as a barrier to particulates, and
its tissue mechanics. Repopulation of the acellular
lung matrix with mixed populations of neonatal
lung epithelial cells resulted in regional-specific
epithelial seeding in correct anatomic locations. To
enhance the survival and differentiation of lung
epithelium, we cultured the matrix in a bioreactor
designed to mimic certain features of the fetal lung
environment, including vascular perfusion and
liquid ventilation (Fig. 1, C and D) (6). Lastly, we
tested the functionality of the engineered lung
tissue by implanting it for short time periods in a
syngeneic rat model (Fig. 1E).

1Department of Biomedical Engineering, Yale University, New
Haven, CT 06520, USA. 2Department of Biomedical Engineer-
ing, Duke University, Durham, NC27708, USA. 3Department of
Anesthesia, Yale University, New Haven, CT 06520, USA. 4De-
partment of Cellular and Molecular Physiology, Yale University,
New Haven, CT 06520, USA. 5Department of Surgery, Yale Uni-
versity, New Haven, CT 06520, USA. 6Department of Internal
Medicine, Yale University, New Haven, CT 06520, USA.

*To whom correspondence should be addressed. E-mail:

Fig. 1. Schema for lung tissue engineering. (A) Native adult rat lung is cannulated in the pulmonary
artery and trachea for infusion of decellularization solutions. (B) Acellular lung matrix is devoid of cells
after 2 to 3 hours of treatment. (C) Acellular matrix is mounted inside a biomimetic bioreactor that allows
seeding of vascular endothelium into the pulmonary artery and pulmonary epithelium into the trachea.
(D) After 4 to 8 days of culture, the engineered lung is removed from the bioreactor and is suitable for
implantation into (E) the syngeneic rat recipient.

30 JULY 2010 VOL 329 SCIENCE www.sciencemag.org538


Preparation of a decellularized lung scaffold.
We harvested lung tissue from adult Fischer 344
rats and treated the lungs in the vascular and airway
compartments with detergent solutions (Fig. 1 and
movie S1). The decellularization solution contains
in a phosphate buffer at 1.0 M salt concentration
(7). Vascular perfusion pressure was maintained
below 20 mmHg, with total time of tissue exposure
typically in the range of 2 to 3 hours. Analysis of
the decellularized lung matrices by means of high-
resolution micro-computed tomography (micro-
CT) revealed intact lung architecture and an intact
arterial tree and microvasculature (Fig. 2, A to C).
The cells and nuclear material were removed with
the decellularization process, but the alveolar sep-
tal architecture remained undisturbed (Fig. 2, E
and F). Fluorometric DNA assay confirmed that

approximately 99% of DNAwas removed by the
decellularization process (Fig. 2G), and immuno-
blotting for major histocompatibility complex I
and II (MHC-I and MHC-II) as well as b-actin
confirmed that decellularized lung matrices were
depleted of these cellular markers (Fig. 2D).
Scanning electron microscopy (SEM) confirmed
that alveolar cells, red blood cells, and cellular
proteins present in native lung were absent in the
decellularized scaffold (Fig. 2, H and I, and fig.
S3). Analysis of the matrix by means of trans-
mission electron microscopy (TEM) revealed that
the ultrastructure of the alveolar septae and the
delicate microvessels surrounding the alveoli re-
mained intact (Fig. 2J and fig. S3). Immuno-
fluorescence and histochemical staining indicated
that collagen, elastin, and laminin were preserved
in decellularized matrices (figs. S4 and S5). Ad-
ditionally, quantitative assays showed that extra-

cellular matrix collagen was preserved, whereas
elastin was partially depleted by the decellulari-
zation process (40% of elastin remains in acel-
lular matrices), and sulfated glycosaminoglycans
(GAGs) were >90% removed (Fig. 2G and fig.
S4). Hence, the lung decellularization protocol
produces an acellular matrix scaffold that retains
the gross, microstructural, and ultrastructural
properties of native lung, yet removal of antigenic
cellular components is essentially complete.

Properties of the lung bioreactor. The lung
bioreactor contains a main reservoir in which the
decellularized lung is mounted, using cannulae
that are inserted into the trachea and pulmonary
artery (fig. S1A). Culture medium is perfused into
the pulmonary artery at physiological pressures
(fig. S1B) (7). To provide negative-pressure ven-
tilation of the engineered lung, a syringe pump
withdraws a defined volume of air from the main
bioreactor to create a negative pressure (Fig. 1C
and figs. S1C and S2). Lung inflation under neg-
ative pressure is accompanied by inhalation of
liquid medium from a trachea reservoir via a
“breathing loop.” For exhalation, the syringe
pump returns air to the main reservoir, causing
lung exhalation of liquid medium into the trachea

To repopulate the decellularized matrix, we
injected cells into the vascular or airway compart-
ments, or both, and maintained the matrix in cul-
ture for up to 8 days (Fig. 1C). This produced an
engineered lung tissue that could then be removed
from the bioreactor for further analysis or that
could be orthotopically transplanted into a rat
recipient (Fig. 1, D and E).

Repopulation of acellular matrix to produce
an engineered lung. Seeding of the decellularized
matrix with mixed populations of neonatal rat
lung epithelial cells (fig. S7) into the airway com-
partment generally resulted in good adherence of
the cells to alveolar structures, as well as to small-
and medium-sized airways (Fig. 3A). Micro-
vascular lung endothelial cells, when injected
into the pulmonary artery of acellular scaffolds,
adhered throughout the scaffold vasculature (Fig.
3B). Seeded pulmonary epithelial cells replicated
rapidly and rarely displayed markers of apoptotic
cell death (Fig. 3, C and D), despite the difficulty
in culturing such cells on standard tissue culture
plastic in vitro (8). This observation suggests that
substrate cues on the acellular lung matrix are
important for pulmonary epithelial cell attach-
ment and replication.

In the biomimetic bioreactor, vascular perfu-
sion greatly enhanced endothelial adhesion and
survival on the lung matrix (fig. S9, A and B).
Negative pressure ventilation had multiple ben-
eficial effects on cultured lung epithelium, in-
cluding enhanced survival of epithelium in distal
alveoli and clearance of epithelial secretions from
the airway tree (fig. S9, C to F). Clearance of
airway secretions through ventilation indicates
that the developing epithelium is in communica-
tion with the airway tree and is not growing ran-
domly within the matrix. In addition, ventilation

Fig. 2. Characterization of acellular lung matrix. (A) 3D micro-CT of the acellular matrix airway
compartment. Large airways are in green. (B) Micro-CT angiography of vascular compartment, thresholded
to visualize only large vessels. In (A) and (B), voxel size is 58 mm; scale bar, 4 mm. (C) Micro-CT angiography
of smaller vessels in acellular lung. Voxel size, 6.5 mm; scale bar, 500 mm. (D) Immunoblot for MHC-1, MHC-
II, and b-actin in native (Nat) and decellularized (Dec) lungs, showing removal of cellular proteins. (E)
Hematoxylin and eosin (H&E) stain of native rat lung. (F) H&E stain of acellular lung matrix. Scale bar, 50
mm in (E) and (F). (G) Collagen (Coll), elastin (Elas), glycosaminoglycan (GAG), and DNA contents of native
lung (black bars) and acellular matrices (hatched bars). Values are mean T SD per lung (n ≥ 4 lungs for all
measures), scaled to 1 for native, with asterisk indicating P < 0.05 for difference between native and acellular matrices. (H) SEM of native rat lung. (I) SEM of acellular matrix. Scale bar, 10 mm in (H) and (I). (J) TEM of acellular lung matrix. Asterisk indicates capillaries in alveolar septa. Scale bar, 5 mm. SCIENCE VOL 329 30 JULY 2010 539


with air—as opposed to with culture medium—
increased the numbers of type I alveolar epithelial
cells, as well as the numbers of ciliated columnar
epithelial cells (fig. S9, G to J). Engineered lungs
also produce pro-surfactant proteins B and C

(pro-SPB, pro-SPC), although we could not de-
tect mature SPB (Fig. 3E). Surfactant proteins are
critical for reducing alveolar surface tension and
enabling lung inflation at physiologically normal

We performed compliance testing on the
engineered lungs under quasi-static conditions by
injecting air into the lungs and monitoring result-
ant pressure changes. Typical compliance curves
for native lung, acellular matrix, and repopulated
engineered lungs are shown in Fig. 3, F to H. The
compliance values were, respectively, 0.35 T 0.08
(n =10measures),0.09T 0.02(n=4measures),and
0.14 T 0.06 (n = 5 measures) mL/mmHg (mean T
SD, P < 0.001 for difference between native and both decellularized and engineered compliances). Compliance values were measured at initial filling of the lungs (arrowheads in Fig. 3, F to H), and thus lower compliances for decellularized matrix and engineered lung mean that these two tissues have higher “opening pressures,” and less functional surfactant, than does native lung. Despite this dif- ference, the overall stress-strain relationships and the ultimate tensile stresses were similar between the three groups (Fig. 3I and fig. S6). Thus, no substantial stiffening or weakening of the extra- cellular matrix occurs in the repopulated, engineered lungs as compared with native adult lungs.

To evaluate the distribution and phenotype of
cells in the engineered lungs, we performed fluo-
rescent immunohistochemical staining (fig. S8).
Endothelial cells seeded into the vasculature were
extensively distributed and expressed CD-31, as
did comparable cells in native rat lung (fig. S8A).
TEM analysis revealed the presence of tight
is consistent with the development of some bar-
rier function (fig. S10). With respect to the seeded
lung epithelium, Clara cell secretory protein
(CCSP)–positive cells, which in native lung are
found in small airways, were found primarily in
very small airway structures after 4 days and in
larger structures after 8 days in engineered lungs
(fig. S8B). Pro-SPC, a marker of type II alveolar
epithelium, is normally present at the vertices of
alveoli in native lung. In the engineered lungs,
pro-SPC expression was diffuse in alveoli and in
small airways after 4 days of culture, but at 8 days
it showed a more native expression pattern at the
vertices of alveoli (fig. S8C). In contrast, aquaporin-
5—a specific marker for type I epithelium—was
found diffusely throughout native alveoli and in
engineered lungs after 4 days but was largely
absent after 8 days (fig. S8D). This observation is
consistent with previous work on neonatal rat
development, which showed that type I cells do
not fully differentiate until the post-natal period,
when air breathing commences (9). Indeed,
engineered lungs that were ventilated with air,
as opposed to liquid culture medium, displayed
more aquaporin-5 expression in alveoli after 8
days of culture (fig. S9, G and H) and also
contained sparse ciliated epithelial cells (fig. S9, I
and J). Additional cell types noted included mes-
enchymal cells and the airway epithelial progen-
itor basal cells (fig. S8, E and F). Hence, the
engineered lungs contained many of the impor-
tant cell types of native lung tissues. In addition,
the spatial distribution of the various cell types
was regional-specific, and with extended culture

Fig. 3. Repopulation of the matrix with lung epithelial and endothelial cells and mechanical assessment
of the engineered lungs. (A) H&E stain of mixed neonatal pulmonary epithelium seeded into acellular
matrix and cultured for 8 days. (B) H&E stain of lung microvascular endothelium seeded into vascular
compartment and cultured for 8 days. Scale bars, 100 mm in (A) and (B). (C) Proliferating cell nuclear
antigen stain of epithelial culture after 8 days; brown nuclei are dividing. (D) Terminal deoxynucleotidyl
transferase–mediated deoxyuridine triphosphate nick end labeling stain of epithelium after 4 days of
culture detects no apoptotic cells (positive cells stain brown; green indicates nuclear counterstain). Scale
bars, 50 mm in (C) and (D). (E) Immunoblots for pro-SPC indicates similar expression for native (Nat) and
engineered (Eng) lung. Engineered lung expresses SPB precursor proteins at 43 and 90 kD, but no mature
SPB. b-actin confirms similar protein loading. (F to H) Quasi-static compliance curves for (F) typical native,
(G) acellular, and (H) engineered lungs; arrows indicate inflation arm of loop. (I) Mean ultimate tensile
strengths (UTS) of native (n = 4), acellular (n = 10), and engineered (n = 5) lungs. Error bars are SD; there
were no significant differences between any groups.

Fig. 4. Implantation of engineered lungs into rats. (A) Tissue-engineered left lung was implanted into
adult Fischer 344 rat recipient and photographed ~30 min later. (B) X-ray image of rat showing the
implanted engineered left lung (white arrow) and the right native lung. (C) H&E stain of explanted
lung. Red blood cells perfusing septa are evident, and some red blood cells are present in airspaces.
Scale bar, 50 mm.

30 JULY 2010 VOL 329 SCIENCE www.sciencemag.org540


periods in the bioreactor the overall pattern of
cellular distribution and differentiation became
more similar to that in native lung tissue.

To determine whether the decellularization
and repopulation methodologies used in our
studies of rat lungs were applicable to human
tissues, we obtained human lung segments from a
tissue bank and treated them with decellularization
solutions for up to 6 hours (7). Histological stain-
ing showed that complete cellular removal was
achieved, with preservation of alveolar architecture
(fig. S11, A and B). We seeded the acellular
matrices with A549 human epithelial carcinoma
cells and endothelial cells derived from human
cord-blood endothelial progenitor cells (7). The
A549 cells adhered well to alveolar surfaces, and
the endothelial cells adhered to the vasculature
(fig. S11, C and D), supporting the notion that
these approaches may also be applicable to hu-
man lung tissues.

Implantation of engineered lungs into rats.
To determine whether engineered rodent lungs
were implantable and functional for gas exchange,
we performed orthotopic left lung transplantation
in four animals (7). Acellular matrices were seeded
with neonatal rat lung epithelium and lung micro-
vascular endothelium and cultured for approximate-
ly 1 week in the bioreactor. For lung implantation,
the native lungs were exposed via left thoracoto-
my, and the left lung was excised. After anticoagula-
tion with heparin, the left mainstem bronchus, left
pulmonary vein, and left pulmonary artery of the
engineered lung were anastomosed to the recip-
ient, the lung was ventilated with 100% oxygen,
and blood flow was reestablished.

In all cases, the engineered lungs were easily
suturable to the recipient and were ventilated with
no visible air leak from the parenchyma (Fig. 4A
and movie S2). All engineered lungs became
perfused with blood over a period of seconds to
minutes, with blood visibly turning from dark to
bright red as the hemoglobin became oxygenated.
Implantation times for engineered lungs ranged
from 45 min to 2 hours. After perfusion and ven-
tilation, blood gas samples were drawn from the
pulmonary artery, left and right pulmonary veins
(veins were clamped and samples drawn from
each lung individually), and the unclamped pul-
monary vein so as to document the extent of gas
exchange occurring in the native and engineered
lungs (Table 1).

Chest x-ray confirmed that the engineered lung
was inflated with air, but the level of inflation was

less than that of the native right lung (Fig. 4B).
Histological evaluation of explanted engineered
lungs revealed red blood cells in large blood
vessels and septal microvessels, and some bleed-
ing into airways, although this was modest (Fig.
4C). Blood gas analysis revealed that the tissue-
engineered lungs were effective in exchanging
oxygen and carbon dioxide (Table 1). Partial pres-
sures of oxygen (PO2) increased from 27 T 7 mmHg
in the pulmonary artery to 283 T 48 mmHg in the
left pulmonary vein, indicating complete hemo-
globin saturation and oxygenation. Hemoglobin
saturation was 100% for both engineered left
lung and native right lung venous samples. In ad-
dition, carbon dioxide removal was efficient, with
CO2 falling from 41 T 13 mmHg in the pulmonary
artery to 11 T 5 mmHg in the left, engineered
pulmonary vein. Although the PO2 in the right
pulmonary vein was higher than in the left (634 T
69 versus 283 T 48 mmHg), this difference may
not be of substantial physiological consequence
because hemoglobin saturation is complete above
oxygen pressures of 100 mmHg (10).

Discussion. To date, cell therapy and tissue
engineering have been applied less extensively to
lung than to other tissues and organs (11–13).
Many efforts in lung regeneration have involved
synthetic scaffolds or simple in vitro culture sys-
tems. Such systems can regenerate certain micro-
scopic features of alveolar architecture but have
not yet produced tissue that can participate in gas
exchange (14, 15). Use of the decellularization para-
digm for respiratory tissue was described in 2008,
when Macchiarini and colleagues implanted a re-
seeded tracheal matrix into a patient with severe
bronchomalacia (16).

In the current work, we have demonstrated the
feasibility of producing an engineered lung that
displays much of the microarchitecture of native
lung and that can effect gas exchange for short
periods of time when implanted into rats. Al-
though these results are encouraging, multiple
issues remain to be addressed before long-term
engineered lung function can be realized. For ex-
ample, alveolar barrier function must be improved
so as to prevent any leakage of blood components
into the airways. This can be accomplished
through iterative improvement in the decellulari-
zation procedure in order to minimize alveolar
septal damage (17). Production of surfactant
should be increased (18), and differentiated co-
lumnar ciliated epithelium should be enhanced by
more prolonged air breathing in culture (19). In

addition, the efficiency of vascular endothelial cov-
erage must be very high throughout the engineered
lung vasculature. This is to prevent exposure of
collagen-containing basement membrane to the
circulation, with consequent thrombosis, because
some clotting was noted at explant. Lastly, a de-
cellularization strategy for lung regeneration will
only become clinically useful when a suitable,
autologous source of pulmonary epithelium can
be identified, such as a resident lung stem cell or
induced pluripotent stem cell (20–24).

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P. Nettesheim, Am. J. Respir. Cell Mol. Biol. 14, 104

20. D. Wang, J. E. Morales, D. G. Calame, J. L. Alcorn,
R. A. Wetsel, Mol. Ther. 18, 625 (2010).

21. J. Yu et al., Science 318, 1917 (2007).
22. C. F. Kim et al., Cell 121, 823 (2005).
23. H. J. Rippon et al., Proc. Am. Thorac. Soc. 5, 717

24. B. Roszell et al., Tissue Eng. Part A 15, 3351 (2009).
25. We thank M. Colehour for help with bioreactor

development and Western blotting, R. Homer for
histologic assessment, and B. Stripp for the CCSP
antibody. This work was funded by Yale University
Department of Anesthesia and by NIH grant HL 098220
(to L.E.N.). T.H.P. is supported by NIH T32 GM007171.
L.E.N. holds stock in Humacyte, a regenerative medicine
company. The authors (L.E.N., T.H.P., and E.A.C.) and
Yale University have filed a patent application related to
tissue engineering of lungs.

Supporting Online Material
Materials and Methods
SOM Text
Figs. S1 to S11
Table S1

9 March 2010; accepted 14 June 2010
Published online 24 June 2010;
Include this information when citing this paper.

Table 1. Gas exchange in engineered lungs implanted into rats. Shown are blood gas values and
oxygen saturations for samples taken from rats implanted with an engineered lung. Values are
mean T SD; n = 3 samples for each sample type except for right pulmonary vein, in which n = 2
samples. Sat, saturation.

Sample location pH PO2 (mmHg) O2 Sat (%) PCO2 (mmHg)

Pulmonary artery 7.30 T 0.06 27 T 7 44 T 20 41 T 13
Right pulmonary vein 7.53 T 0.08 634 T 69 100 T 0 20 T 1
Left (implant) pulmonary vein 7.68 T 0.28 283 T 48 100 T 0 11 T 5
Mixed pulmonary veins 7.58 T 0.08 495 T 174 100 T 0 18 T 3 SCIENCE VOL 329 30 JULY 2010 541



994 Vol 380 September 15, 2012

Lancet 2012; 380: 994–1000

Published Online
July 26, 2012

See Comment page 957

See Online for a video interview
with Martin Birchall

Department of Cardiothoracic
Surgery (Prof M J Elliott MD,

S Speggiorin MD, A Fierens RNC,
D Vondrys MD), Department of

Surgery (P De Coppi MD),
Department of Radiology

(D Roebuck MB, C McLaren DCR),
and Ear Nose and Throat
Surgery (L Cochrane MD,

C Jephson FRCS), Great Ormond
Street, Hospital for Children,

London, UK; Centre for
Respiratory Research

(C R Butler MRCS, S Janes MD),
Centre for Nanotechnology
and Regenerative Medicine

(C Crowley MSc,
Prof A M Seifalian PhD), and

Centre for Molecular Cell
Biology (N J Beaumont PhD,

Prof J J Hsuan PhD), University
College London, London, UK;
Paul O’Gorman Laboratory of

Cellular Therapeutics,
Department of Haematology,

Royal Free Hospital, London,
UK (E Samuel MSc,

M W Lowdell PhD); School of
Veterinary Sciences, University

of Bristol, Bristol, UK
(T Cogan PhD); Department of

Cell Techniques and Applied
Stem Cell Biology, University of

Leipzig, Leipzig, Germany
(Prof A Bader MD); and

University College London Ear
Institute, The Royal National

Throat, Nose and Ear Hospital,
London, UK

(Prof M A Birchall MD)

Correspondence to:
Prof Martin A Birchall, University

College London Ear Institute,
Royal National Throat Nose and
Ear Hospital, 330332, Gray’s Inn

Road, London WC1X 8EE, UK

Stem-cell-based, tissue engineered tracheal replacement in
a child: a 2-year follow-up study
Martin J Elliott, Paolo De Coppi, Simone Speggiorin, Derek Roebuck, Colin R Butler, Edward Samuel, Claire Crowley, Clare McLaren, Anja Fierens,
David Vondrys, Lesley Cochrane, Christopher Jephson, Samuel Janes, Nicholas J Beaumont, Tristan Cogan, Augustinus Bader,
Alexander M Seifalian, J Justin Hsuan, Mark W Lowdell, Martin A Birchall

Background Stem-cell-based, tissue engineered transplants might off er new therapeutic options for patients, including
children, with failing organs. The reported replacement of an adult airway using stem cells on a biological scaff old
with good results at 6 months supports this view. We describe the case of a child who received a stem-cell-based
tracheal replacement and report fi ndings after 2 years of follow-up.

Methods A 12-year-old boy was born with long-segment congenital tracheal stenosis and pulmonary sling. His airway
had been maintained by metal stents, but, after failure, a cadaveric donor tracheal scaff old was decellularised. After a
short course of granulocyte colony stimulating factor, bone marrow mesenchymal stem cells were retrieved
preoperatively and seeded onto the scaff old, with patches of autologous epithelium. Topical human recombinant
erythropoietin was applied to encourage angiogenesis, and transforming growth factor β to support chondrogenesis.
Intravenous human recombinant erythropoietin was continued postoperatively. Outcomes were survival, morbidity,
endoscopic appearance, cytology and proteomics of brushings, and peripheral blood counts.

Findings The graft revascularised within 1 week after surgery. A strong neutrophil response was noted locally for the
fi rst 8 weeks after surgery, which generated luminal DNA neutrophil extracellular traps. Cytological evidence of
restoration of the epithelium was not evident until 1 year. The graft did not have biomechanical strength focally
until 18 months, but the patient has not needed any medical intervention since then. 18 months after surgery, he
had a normal chest CT scan and ventilation-perfusion scan and had grown 11 cm in height since the operation. At
2 years follow-up, he had a functional airway and had returned to school.

Interpretation Follow-up of the fi rst paediatric, stem-cell-based, tissue-engineered transplant shows potential for this
technology but also highlights the need for further research.

Funding Great Ormond Street Hospital NHS Trust, The Royal Free Hampstead NHS Trust, University College
Hospital NHS Foundation Trust, and Region of Tuscany.

There are no universally eff ective solutions for the
treatment of advanced structural disorders of the large
airways in children. Such children need frequent stays in
hospital. Although slide tracheoplasty is the primary
treatment of choice for some children, others develop
recurrent stenoses.1 Stent erosion and death can occur.2
Fetuses with laryngotracheal agenesis or severe stenosis
identifi ed before birth might be aborted because these
abnormalities are regarded as fatal.3

In the past decade, tissue-engineered structures re-
populated with cells or stem cells have been used clinically.
Atala and colleagues4 used collagen scaff olds reseeded
with urothelial and muscle cells to repair bladder defects
in patients with myelo meningocele, and the successful
clinical application of a stem-cell-based tracheal replace-
ment in a woman with end-stage airway disease with
6 months follow-up has been reported.5 In 2011, good
short-term (5 months) outcomes were reported in a
patient who received a similar stem-cell-based tracheal
graft, but in this study a nanocomposite trachea was used
as the scaff old.6 However, the long-term outcomes of these

and other patients who have received such grafts on
compassionate grounds have yet to be published pending
adequate follow-up, and there is no previous report of
outcomes after a tracheal graft in a child.

The potential clinical advantages of autologous stem-
cell-derived transplants are that patients who receive
them would not need immunosuppression and that the
transplants are hypothesised to be remodelled by local
stroma to simulate native tissue.5 By contrast, allo-
trans plantation is associated with signifi cant long-term
mortality due to infection and immunosuppression,
especially in the respiratory system.7 There is also a
paucity of donors for transplantation. Thus, there is a
signifi cant unmet need for novel methods of replacing
and regenerating human tissue.

The ideal endpoints for tracheal replacement in children
are normal airway and lung function, appropriate growth,
high quality of life, and the elimination of the need for
repeated surgical inter ventions. Here, we describe the
case of a child who received a stem-cell-based tracheal
replacement as an urgent compassionate-use procedure
and report fi ndings after 2 years of follow-up.

Articles Vol 380 September 15, 2012 995

The recipient
A child born with long-segment congenital tracheal sten-
osis and pulmonary sling underwent autologous patch
tracheoplasty at Great Ormond Street Hospital NHS Trust
(London, UK) at 6 days old. He could not be extubated after
surgery because of collapse and scarring of the patch and
severe bilateral bronchomalacia. Balloon-expandable stain-
less steel stents (Palmaz, Cordis, Miami Lakes, FL, USA)
were implanted, with marked clinical improvement.

At 3 years old, the child had substantial bleeding into his
airway. Emergency bronchoscopy and CT angi o graphy
revealed erosion by the stent into the aorta. Emergency
aortic repair was done with a bovine pericardial patch
(Synovis, St Paul, MN, USA), and the impacted stents and
trachea were excised and replaced by a tracheal homo-
graft.8 The homograft was replaced 1 week later by another
(stented) homograft after mediastinitis occurred. After
3 months in hospital, the patient made an excellent re-
covery. Over the ensuing years, he needed occasional inter-
ventions, including further stents for recurrent stenosis.

At 10 years old, the patient suff ered a second haemor-
rhage. Findings from bronchoscopy and CT scan suggested
erosion of tracheal stents, creating a new aortotracheal
fi stula. Bleeding stopped spontaneously, which provided
us with time to plan urgent reconstruction.

Use of tracheal homografts had been discontinued
and so other options were discussed. Tracheal allo grafting9
was dismissed due to the prospect of lifelong immuno-
suppression. In view of our previous success with an
autologous stem-cell-based tracheal replacement, the
child’s parents were approached and asked to consider the
use of a similar method for their child. The emergent
nature of his disorder, unlike that of the adult recipient,
meant that a more direct protocol for graft preparation was
needed, and so the published technique was adapted using
methods previously applied success fully to bone, skin, and
nerve regeneration.10 After approval by the Medicines and
Healthcare Products Regulatory Agency (MHRA) and the
institutional Clinical Ethics Committee, his parents con-
sented to the procedure. An appropriate scaff old was
sought and the patient was prepared for surgery.

Pretransplant preparation
The patient received 10 mg/kg granulocyte colony
stimulating factor (G-CSF; Chugai, London, UK) daily
for 3 days before surgery to mobilise haemopoietic stem
cells and endothelial progenitors11 and induce mesen-
chymal stem cell (MSC) proliferation.12 We measured
leucocyte counts daily.

After general anaesthesia, 50 mL autologous bone
marrow was aspirated into sterile, heparinised tubes,
diluted 1:1, and mononuclear cells were isolated by
discontinuous density gradient separation. Sterility test-
ing was done on the washings. Cells were re-suspended,
topped up, and counted. The total mononuclear cell count
was 2·56×10⁸, of which 5·51×10⁶ were CD34+ or CD45wk

haematopoietic stem cells and 1·68×10⁶ were CD73+,
CD90+, CD105+, CD117+, or CD45+ MSCs. The cell
suspension was transferred into a 60 mL Cryocyte bag
(Miltenyi, Bisley, UK) supplemented with 10 IU protamin
(Wockhardt, Wrexham, UK) and shipped in a tem-
perature-monitored container (4°C) to the operating room
at Great Ormond Street Hospital NHS Trust.

A CT scan was done to identify the dimensions needed
for the scaff old. An allogeneic trachea of appropriate size
was retrieved by permission of the Tuscany regional
authorities from a 30-year-old female donor. Infectious
disease markers were negative. The scaff old was pre pared
in the Regenerative Surgery Laboratories of the University
Hospital Careggi, Florence, Italy, using a published
protocol,13 and was released by quality control after con-
fi rmation of sterility and absence of HLA-I+ cells. 3 days
before surgery, the scaff old was imported (4°C) in
phosphate-buff ered saline supplemented with penicillin,
streptomycin, and amphotericin B, in compliance with
UK Human Tissue Authority codes and local licence
number 11016. 10 000 IU erythropoietin (Roche, Welwyn,
UK), 200 IU G-CSF (Neupogen, Cambridge, UK) and
50 μg transforming growth factor β (TGFβ; R&D Systems,
Abingdon, UK) were contained in separate syringes and
transported with the scaff old and cells.

Tracheal replacement surgery
During surgery, the head-down tilt position, cardio-
pulmonary bypass, and progressive cooling to 18°C were
used. With the heart decompressed, a resternotomy was
done. Use of right atrium and superior vena cava venous
lines permitted cardiac isolation and great vessels
were mobilised. At 18°C, the aorta was cross-clamped,
anterograde cardioplegia was instilled, head vessels were
snared, and the circulation was stopped. After dissection,
the stent that was entering the aorta was identifi ed, as
were others buried in the tracheal wall (fi gure 1A). The
aortic defect was repaired with bovine pericardium.
Circulation was resumed and re-warming commenced.

The trachea was transected above the upper metal stent
and below the lower metal stent, leaving a 7 cm gap be-
tween the upper trachea and the carina (fi gure 1B). Patches
of tracheal epithelium were removed from the excised
trachea, cut into stamp grafts and retained. The stents in
the main bronchi were trimmed back to provide cuff s of
unstented bronchi for anastomosis. Bronchi were dilated
with an 8 mm Hegar dilator (Lyall Willis, Hastings, UK).

In the operating room, the scaff old was saturated with
the cell suspension. The mucosal stamps were placed as
free grafts at regular intervals within the lumen. An
absorbable polydioxanone (PDO) tracheal stent (Ella-Cs,
Hradec Kralove, Czech Republic) measuring 12×72 mm
was sutured in place (5/0 PDS, II, Ethicon, Edinburgh,
UK). The construct was saturated with human recom-
binant erythropoietin (hrEPO) and G-CSF, and TGFβ was
injected into the tracheal rings (fi gure 1D) to increase
angiogenesis, improve autologous MSC recruit ment, and


996 Vol 380 September 15, 2012

induce chondrocyte diff erentiation. The construct was
anasto mosed superiorly and inferiorly using horizontal
mattress interrupted sutures (4-0 PDS II, fi gure 1E). Before
completing the anastomoses, a new trans-nasal endo-
tracheal tube was placed under direct vision. Two small air
leaks were sealed and the child was weaned from
cardiopulmonary bypass. The omentum was mobilised
and interposed between the trachea and heart to reduce
the possibility of future fi stulae and increase graft
vascularity (fi gure 1F). On alternate days after surgery,

hrEPO (10 000 IU for 2 weeks) and G-CSF (10 mg/kg for
1 week) were administered.

Role of the funding source
The sponsors of the study had no role in the study design,
data collection, data analysis, data interpretation, or
writing of the report. The corresponding author had full
access to all the data in the study and had fi nal
responsibility for the decision to submit for publication.

Within 3 h of the surgery, ventilation became problematic
and bilateral air trapping occurred. Fibreoptic bron-
choscopy showed substantial narrowing of the origin of
both bronchi due to the longitudinal rigidity of the
absorbable stent. A temporary stent (Niti-S, Taewoong,
Seoul, South Korea) was implanted in each bronchus
under fl uoroscopy, which resulted in an immediate
improvement in ventilation. These stents were removed
before extubation on day 26 after surgery. There was an
initial increase in the number of circulating leucocytes
(31·1×10⁹/L [SD 6·6×10⁹/L]) between days 2 and 8 after
surgery, which corresponded to the period of application
of G-CSF and hrEPO. This period was also the only time
when circulating CD34+ cells (0·71×10⁹/L [SD 0·05×10⁹/L])
could be detected. Leucocyte counts normalised from
day 9 (9·14 ×10⁹/L [SD 1·18×10⁹/L]; appendix).

There was bleeding on contact from the internal
lumen of the graft by 1 week after surgery, which proved
that angiogenesis was occurring. The patient needed
regular bronchoscopy for removal of dense secretions
for 8 weeks (fi gure 2A). Assessment of the secretions
showed that they included no cells, had a high DNA
content, and had a net-like microscopic appearance. The
features were identifi ed as those of DNA neutrophil
extracellular traps (NETs).14,15 SDS gel frac tionation,
tryptic digestion, and nano-liquid chroma tography mass
spectrometry identifi ed a protein profi le consistent with
this diagnosis (fi gure 3).14 The secretions were treated
with a combination of DNase and physiotherapy, and
cleared as epithelialisation pro gressed. The patient was
discharged on day 63.

6 weeks after surgery the stent had dissolved and there
was mild collapse of the proximal graft. A shorter
(10×45 mm) PDO stent was implanted under fl uoroscopy.
The patient underwent bronchoscopy or balloon dilatation
under fl uoroscopy, or both, regularly for 6 months
(fi gure 2B). The major reason for further bronchoscopy or
balloon dilation was mucus retention and crusting within
the native bronchi in which there were still embedded
metal stents. At 5 months, after dissolution of the latest
stent, we remained concerned about the rigidity of the
proximal graft, and so overlapping, self-expanding Nitinol
stents (S.M.A.R.T. Control, Cordis, Waterloo, Belgium)
were implanted into the trachea. At 6 months after the
initial surgery the graft seemed stable, the patient’s airway
was patent, and he returned to school.



Figure 1: Surgical procedure
(A) During surgery the airway was found to be severely stenotic with multiple stents including one entering the
ascending aorta. (B and C) The old homograft trachea was removed and replaced by the engineered graft.
(C) The aortic defect was closed with a bovine pericardial patch and air leaks sealed. (D) Transforming growth
factor β was injected into tracheal rings in the operating theatre before (E) implantation of the recellularised
graft. (F) Before closing, an omental wrap was brought up to cover the graft. The graft sits in the anatomical
position to the right of the ascending aorta.


Figure 2: Bronchoscopic appearances
(A) Microlaryngobronchoscopy 15 days after the transplant showing a dense web covering the stent and partially
occluding the lumen (A), which was cleared by regular bronchoscopies and DNAase. (B) Image at 6 months,
showing that reabsorption of the stent (white areas) caused so-called cobblestones of granulation tissue with little
normal epithelium. (C) At 15 months after surgery, the graft seemed to be patent, with healthy mucosa.

See Online for appendix

Articles Vol 380 September 15, 2012 997

The patient’s last endoscopy (15 months after surgery)
showed complete epithelialisation (fi gure 2C), and
cytology of tracheal brushings showed healthy, ciliated
respiratory epithelial cells (fi gure 4D). At 18 months, he
had his last fl uoroscopic balloon dilatation because the
malacic seg ment had strengthened such that he had not
needed any further admissions to hospital. As of May 7,
2012, he was well, active, and had grown 11 cm and 5 kg
since graft im plantation. His lungs appeared normal on
CT scan, without bronchiectasis or air trapping, and a
ventilation-perfusion scan at 12 months was normal
(fi gure 5). As of May 13, 2012, there has been no
serological or clinical evidence of rejection of the graft
and a comprehensive screen of his serum at 15 months
showed no anti-HLA antibodies.

Histological assessment of the homograft trachea
removed at the time of surgery showed an infl amed
mucosa overlying dense fi brous tissue and islands of
cartilage unlike normal tracheal architecture (fi gure 4B).
Histology of the decellularised scaff old showed complete
decellularisation with good retention of tracheal archi-
tecture (fi gure 4C) and absence of MHC expression (not
shown). High-resolution proteomic analysis of the scaff old
by ion-trap mass spectrometry identifi ed 166 proteins,
including several extracellular matrix components. Bio-
informatic analysis (IPA, Ingenuity Systems, Redwood
City, CA, USA) identifi ed a broad range of potential bio-
logical roles for these proteins.

We report a stem-cell-based tissue replacement in a
child and long-term follow-up of a stem-cell-based
tissue-engineered graft (panel). The child is well,
growing, and had not needed medical intervention for
6 months by May 5, 2012.

Because the protocol used in this study was devised in an
emergency, we applied empirically a new combination of
technologies on the basis of previous clinical successes in
non-airway settings (ie, bone, skin, and nerve regeneration).


Description Mw (Da) PLGS score Peptides Coverage (%)
















Alpha 2 macroglobulin

Complement C3


Ig alpha 1 chain C region

Ig alpha 2 chain C region

Ig gamma 2 chain C region

Ig kappa chain C region

Ig kappa chain V III region SIE

Ig kappa chain V III region WOL

Ig mu chain C region


Polymeric immunoglobulin receptor


Serum albumin


163 188

187 029

51 643

37 630

36 503

35 877

11 601

11 767

11 738

49 275

78 131

83 231

76 999

69 321

141 059




































Figure 3: Identifi cation of protein in the tracheal exudate
Proteins in the tracheal exudate identifi ed in the early weeks (sampled postoperative week 2) were separated using SDS-PAGE and stained with colloidal Coomassie
Blue (A). Destained gel slices were digested with trypsin (Promega, Southampton, UK), fractionated by high-performance liquid chromatography (NanoAcquity,
Waters, Manchester, UK), and analysed using an in-line Q-TOF mass spectrometer (Waters). (B) The table shows the proteins identifi ed from at least two peptides and
with a PLGS score greater than 10. PLGS=Protein Lynx Global Server.



250 μm

100 μm 20 μm

100 μm

Figure 4: Findings on cytology
Haematoxylin and eosin staining of (A) normal trachea compared with (B) the patient’s previous tracheal homograft
removed at the time of surgery, which shows an epithelialised lining but atypical gland formation. (C) A sample of the
decellularised tracheal graft used in this study shows loss of cells but preservation of normal architecture. (D) Bronchial
brushing taken from the middle of the graft 1 year after surgery shows a cluster of ciliated cells.


998 Vol 380 September 15, 2012

Thus, to minimise delays, there was no previous expansion
of epithelial cells and MSCs, nor any chondrocytic
diff erentiation of MSCs.5 Instead, we used an intraoperative
protocol, which was similar to those used in clinical trials
of MSCs for patients with myocardial infarction.17 Not
undertaking long-term culture of MSCs also has the
potential advantage of avoiding the risk of malignancy.18
We aimed to create an in-vivo microenvironment that
represented some of the events that occur during the
normal physiological response to injury. A similar method
is in phase 2 clinical trials of bone, skin, and nerve
regeneration.10 We hypothesise that this altered protocol, in
addition to the length of the graft, the presence of an
absorbable stent, and the underlying diff erent physiology
and regenerative potential of children’s compared with
adults’ tissues, were responsible for the diff erences in
clinical course and outcomes from the published adult
case, at least at the 6-month timepoint.5 Specifi cally, the
graft in the present study took longer to epithelialise and
did not have proximal rigidity until almost 2 years.
However, at last follow-up the boy was alive, growing, had
normal lung function, and had returned to school.

A key criterion for paediatric implants is that of
growth potential. In this study, although we were
unable to measure graft length, there was no CT
evidence of shortening of the graft, as has been
previously described, for example, with an alternative
aortic allo graft approach.19 At age 13 years the child’s
torso is not expected to elongate much further as his
height increases and so the growth demands on this
graft are limited. However, experimental evidence of
graft growth is crucial for the clinical use of similar
protocols of transplantation for children of all ages.
Equally crucial is rapid vascularisation. As with the
adult case,5 touch bleeding on the internal graft surface
was visible by 1 week, which proved that rapid angio-
genesis was occurring.

In both this study and a previous case,5 a cadaveric
donor trachea was decellularised, with successful removal
of cellular components including MHCs. Neither patient
had developed rejection by May, 2012, and the child had
not developed anti-donor antibodies by 20 months. These
fi ndings, in addition to reports of preclinical success with
similar methodologies for heart and lung grafts,20,21
suggest that decellularised scaff old-based technologies
could be an immuno suppression-free alternative to
conventional transplantation.

In the UK, patients operated upon under a Hospitals
Exemption Certifi cate on compassionate grounds, as was
the case with the patient in this study, are not treated as
research patients. Thus, we did not label the applied cells
and so cannot comment on whether the eventual stromal
and epithelial cells originated from those implanted or
from cells recruited from neighbouring tissues. Future
preclinical and clinical trials should incorporate markers
that will answer the question of the exact contribution of
applied cells to the fi nal result.

Many clinicians assume that decellularised scaff olds
are inert composites of structural proteins. Proteomic
measurement of non-structural, or minor structural,
proteins has been diffi cult because of the dominance of
collagen and elastin in protein preparations. In this
study, with new techniques we identifi ed 166 proteins
with diverse functions relevant to regenerative medicine
(eg, angiogenesis and immunity) that were preserved
despite decellularisation, although no MHC molecules
were found. We hypothesise that many of these pro teins
are crucial to revascularisation, cell migration, and
diff erentiation in tissue-engineered organs and repre-
sent a major diff erence from synthetic scaff olds. There-
fore, proteomic analysis might be a valuable addition to
release criteria for biological scaff olds.

TGFβ was added to the scaff old to induce chondrocytic
diff erentiation, G-CSF to boost autologous MSC recruit-
ment, and hrEPO to increase angiogenesis. G-CSF is
used to mobilise bone marrow progenitor production
before haemopoietic cell transplantation.22 Although
some studies report a benefi cial eff ect of G-CSF on MSC
mobilisation,12,23 others suggest the opposite eff ect.17




Figure 5: Follow-up scans
(A) CT axial scan and (B) coronal scan done 12 months after surgery show the tracheal graft (arrows) surrounded by
omental fat (*). The lumen of the graft is narrow (6 mm) and its wall is thick (3–4 mm). Growth in length of the
graft was not seen on serial images, possibly because growth in height of the child was not matched by lengthening
of the chest. (C) A lung scan (ventilation-perfusion) at 18 months showed normal bilateral ventilation (the left lung
is contributing 45% to the total ventilation and the right lung 55%). There is a slight reduction in perfusion in the
left lung (receiving 37% of the right heart output) compared with the right lung (63%).

Articles Vol 380 September 15, 2012 999

Identifi cation of the contribution of G-CSF to the survival
and function of the graft in one patient in the short and
long term is not possible. However, we hypothesise that
system ic application of G-CSF increased leucocyte
counts in week 1 and contributed to NET accumulation
in the trachea in the fi rst 6 weeks after surgery.23,24

hrEPO is used clinically to support erythropoiesis in
patients with cancer and renal disease.25 Pretreatment
with erythropoietin might improve the survival of cells
within tissue where angiogenesis is not yet adequate to
fully support respiration, by a mechanism mediated by
nitric oxide and vascular endothelial growth factor.26
Angiogenesis, measured by appearance, contact bleed-
ing, and laser doppler fl uxmetry, was equally fast in the
previously reported adult patient5 as in the child in the
present study. Despite the substantial increase in graft
mass in the child, we can only speculate about the added
angiogenic eff ect of hrEPO. Findings are further
confounded by the use of an omental fl ap, because the
purpose of it is to provide an improved vascular bed for
the graft. More research into angiogenic mechanisms in
re cellularised regenerative constructs is needed.

We hypothesised that TGFβ, a key signal for chon-
drocytic diff erentiation of MSCs,27 would enable repopu-
lation of the preserved scaff old cartilage niche and provide
adequate biomech anical support in the long term.
However, TGFβ is also a powerful promoter of myo-
fi broblasts and scar tissue28 and restricts epithelial cell
survival and migration,29 both of which are undesirable
actions during the regeneration of tissue-engineered
trachea. The absence of rigidity in the proximal trachea
suggests that TGFβ did not support adequate cartilage
regeneration throughout the graft, although the length of
the graft, presence of PDO stents, or the absence of a
preoperative chondrocytic diff eren tiation step in the
process5 might also have been responsible.

PDO stents30,31 have been used in six lung transplant
patients who needed multiple insertions,31 which was
also the case with the child in this study. All six patients
were free of stenosis at a median of 24 months
(range 7–44). Recent experience in children with airway
stenosis is similar.32 The PDO stents were quick to apply
and provided circumferential support for 8 weeks, but
the absence of vertical elasticity was a problem and they
might have contributed to NET formation.

Analysis suggested that the problematic tracheal exu-
date in this study was DNA NETs.14,15 The macroscopic
appearance of NETs is poorly described in man. Their
perceived role is to prevent bacterial colonisation and
dissemination, but formation can cause tissue damage.15
We hypothesise that neutrophil recruitment induced by
the graft and stent plus G-CSF treatment were causative,
and that NET resolution parallels the development of
new epithelium.

The epithelium was patchy by 2 months, although stents
caused discontinuity. The presence of viable, ciliated
epithelial cells was confi rmed on cytology at 1 year, when

mucosal continuity was noted throughout. Epithelialisation
occurred later than in the previous adult case,5 where
mucosal coverage was achieved at 1 month and mucociliary
clearance by 6 months.5 The need for early mucosal
coverage and mucociliary clearance for airway grafts in
patients, many of whom have compromised bronchial or
lung function, or both, means that research into
mechanisms of regeneration of the respiratory mucosa is
crucial, as is identifi cation of key stem or progenitor cells
and migration and diff erentiation factors.

This report should be compared with other published
case reports of tissue-engineered airways, described briefl y
earlier5,6,13 and reviewed in more detail elsewhere.3 Sub-
stantial areas for improvement in outcomes were identifi ed
by this experience of a stem-cell-based, paediatric tracheal
replacement; specifi cally, the need for biomechanical
strength throughout the graft and speedy, effi cient
restoration of the mucosa. The response of children to
implants will probably diff er from that of adults in
important ways, including the need to accommodate
growth. Urgent research is needed to convert one-off ,
compassionate-use suc cesses, such as the one described in
this study, into more widely applicable clinical treatments
for the thousands of children with tracheal stenosis and
malacia worldwide.
MJE, SS, and others did the transplant surgery. Postoperative endoscopies
and general medical care were done by MJE, SS, AF, DV, LC, CJ, and
MAB. Radiology, radiologically-guided procedures, and stent placements
were done by DR, CM, DV, and MAB. Good Manufacturing Practice cell
and cytokine preparation was done by ES, CC, and MWL. Analysis of
cytology was done by CB and SJ; histology by CB, CC, TC, and AMS; and
proteomics by CB, NJB, and JJH. Advice on clinical use of cytokines was
provided by AB. MJE, MWL, and MAB designed the protocol. Further data
collection was done by PDC, CB, and DV. MJE, PDC, ES, MWL, and MAB
did the literature search, data interpretation, and writing of the report.

Confl icts of interest
We declare that we have no confl icts of interest.

This work was supported by Great Ormond Street Hospital NHS Trust, The
Royal Free Hampstead NHS Trust, and University College Hospital NHS
Foundation Trust (all London, England), and by a grant

Panel: Research in context

Systematic review
We searched PubMed for all publications, including clinical trials, meta-analyses, and
reviews, with the terms “graft” and “short-term” or “long-term”. However, we did not
restrict our searches to only papers that included the phrase “stem cells” and identifi ed
only two similar case reports, both in adults and with 6 and 5 months’ follow-up
respectively.5,6 Although several conventional treatments are available for the treatment
of congenital tracheal stenosis, no proven treatments exist for patients with end-stage
disease.1,16 Evidence from studies in animals suggests that stem-cell-based tissue
engineered tracheal implants could be useful as part of new treatment strategies for
incurable tracheal stenosis or malacia in children.

This study describes a stem-cell-based organ transplant in a child and is the fi rst in either
adults or children to report long-term follow-up (2 years).


1000 Vol 380 September 15, 2012

(pd 239- 28/04/2009, delib. GRT 1210/08) from the Region of Tuscany
(Italy) entitled “Clinical laboratory for complex thoracic respiratory and
vascular diseases and alternatives to pulmonary transplantation”. Both
Great Ormond Street Hospital and University College Hospital receive
translational research funding from the UK Department of Health’s
National Institute for Health Research Biomedical Research Centres
scheme (MJE, PDC, SJ, and MAB). MJE is Director of the Service for
Severe Tracheal Disease in Children, funded by the National Health Service
National Commissioning Mechanisms. SJ is a recipient of a Wellcome
Trust Senior Fellowship in Clinical Science. Some laboratory work was
supported by a Medical Research Council Translational Stem Cell Research
Committee grant to MB (G1001539) and a Great Ormond Street Hospital
Charity grant to PDC. We thank Caroline Doyle, whose superb
administrative skills were essential to the co-ordination of this procedure.
We also thank the anaesthetic, technical, paramedical, and nursing staff in
operating theatres, intensive care units, and wards at Great Ormond Street
Hospital for Children National Health Service Trust, as well as the senior
management of the Trust who approved the internal funding. We thank
staff at the Royal Victoria Hospital NHS Trust in Belfast who saved the
child’s life at fi rst presentation and cared for him on several occasions over
the past 11 years. All the staff of the Paul O’Gorman Laboratories for Cell
Therapy at the Royal Free Hospital Hampstead NHS Trust contributed to
cell, cytokine, and graft preparation, as did many members of the
Nanotechnology and Surgical Sciences Laboratories at the same institution.
We thank our patient’s parents, who were an essential and supportive part
of the team and decision-making processes. We also thank the Tuscany
Transplant Authority, the Thoracic Surgeons, General and Medical
Directors of the University Hospital Careggi in Florence (Italy), staff at the
UK’s Medicines and Healthcare Products Regulatory Agency (MHRA),
especially Ian Rees, who gave timely and free advice on regulatory aspects
of this case. The Chairman and members of the Clinical Ethics Committee
at Great Ormond Street Hospital constructively considered all aspects to the
case and helped with design of the parent’s information sheet. We fi nally
pay particular tribute to the courageous and inspiring young man himself.

1 Kocyildirim E, Kanani M, Roebuck D, et al. Long-segment tracheal

stenosis: slide tracheoplasty and a multidisciplinary approach
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2 Jacobs JP, Quintessenza JA, Botero LM, et al. The role of airway
stents in the management of pediatric tracheal, carinal and
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3 Lange P, Fishman JM, Elliott MJ, De Coppi P, Birchall MA. What
can regenerative medicine off er for infants with laryngotracheal
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4 Atala A, Bauer SB, Soker S, Yoo JJ, Retik AB. Tissue-engineered
autologous bladders for patients needing cystoplasty. Lancet 2006;
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5 Macchiarini P, Jungebluth P, Go T, et al. Clinical transplantation of
a tissue-engineered airway. Lancet 2008; 372: 2023–30.

6 Jungebluth P, Evren A, Baiguera S, et al. Tracheobronchial
transplantation with a stem-cell-seeded bioartifi cal nanocomposite:
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7 US Department of Health and Human Services. 2009 Annual
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8 Jacobs JP, Quintessenza JA, Andrews T, et al. Tracheal allograft
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9 Delaere P, Vranckx J, Verleden G, De Leyn P, Van Raemdonck D.
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10 Bader A, Lorenz K, Richter A, et al. Interactive role of trauma
cytokines and erythropoietin and their therapeutic potential for
acute and chronic wounds. Rejuvenation Res 2011; 14: 57–66.

11 Engelmann MG, Theiss HD, Hennig-Theiss C, et al. Autologous
bone marrow stem cell mobilization induced by granulocyte
colony-stimulating factor after subacute ST-segment elevation

myocardial infarction undergoing late revascularization: fi nal results
from the G-CSF-STEMI (Granulocyte Colony-Stimulating Factor
ST-Segment Elevation Myocardial Infarction) trial. J Am Coll Cardiol
2006; 48: 1712–21.

12 Brouard N, Driessen R, Short B, Simmons PJ. G-CSF increases
mesenchymal precursor cell numbers in the bone marrow via an
indirect mechanism involving osteoclast-mediated bone resorption.
Stem Cell Res 2010; 5: 65–75.

13 Baiguera S, Jungebluth P, Burns A, et al. Tissue engineered human
tracheas for in vivo implantation. Biomaterials 2010; 31: 8931–38.

14 Brinkmann V, Reichard U, Goosmann C, et al. Neutrophil
extracellular traps kill bacteria. Science 2004; 303: 1532–35.

15 Lögters T, Margraf S, Altrichter J, et al. The clinical value of neutrophil
extracellular traps. Med Microbiol Immunol 2009; 198: 211–19.

16 Elliott M, Roebuck D, Noctor C, et al. The management of
congenital tracheal stenosis. Int J Pediatr Otorhinolaryngol 2003;
67 (suppl 1): S183–92.

17 Ripa RS, Haack-Sørensen M, Wang Y, et al. Bone marrow derived
mesenchymal cell mobilization by granulocyte-colony stimulating
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18 Jeong JO, Han JW, Kim JM, et al. Malignant tumor formation after
transplantation of short-term cultured bone marrow mesenchymal
stem cells in experimental myocardial infarction and diabetic
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19 Wurtz A, Porte H, Conti M, et al. Surgical technique and results of
trachea and carinal replacement with aortic allografts for salivary
gland-type carcinoma. J Thorac Surg 2010; 140: 387–93.

20 Ott HC, Matthiesen TS, Goh SK, et al. Perfusion-decellularized
matrix: using nature’s platform to engineer a bioartifi cial heart.
Nat Med 2008; 14: 213–21.

21 Petersen TH, Calle EA, Zhao L, et al. Tissue-engineered lungs for in
vivo implantation. Science 2010; 329: 538–41.

22 Siddiq S, Pamphilon D, Brunskill S, Doree C, Hyde C, Stanworth S.
Bone marrow harvest versus peripheral stem cell collection for
haemopoietic stem cell donation in healthy donors.
Cochrane Database Syst Rev 2009; 1: CD006406.

23 Tatsumi K, Otani H, Sato D, et al. Granulocyte-colony stimulating
factor increases donor mesenchymal stem cells in bone marrow
and their mobilization into peripheral circulation but does not
repair dystrophic heart after bone marrow transplantation. Circ J
2008; 72: 1351–58.

24 Kassis I, Zangi L, Rivkin R, et al. Isolation of mesenchymal stem
cells from G-CSF-mobilized human peripheral blood using fi brin
microbeads. Bone Marrow Transplant 2006; 37: 967–76.

25 Rizzo JD, Brouwers M, Hurley P, et al, for the American Society of
Hematology and the American Society of Clinical Oncology Practice
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American Society of Clinical Oncology clinical practice guideline
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cancer. Blood 2010; 116: 4045–59.

26 Rezaeian F, Wettstein R, Amon M, et al. Erythropoietin protects
critically perfused fl ap tissue. Ann Surg 2008; 248: 919–29.

27 Moretti M, Wendt D, Dickinson SC, et al. Eff ects of in vitro
preculture on in vivo development of human engineered cartilage
in an ectopic model. Tissue Eng 2005; 11: 1421–28.

28 Antoshina E, Ostrowski LE. TGF beta 1 induces growth arrest and
apoptosis but not ciliated cell diff erentiation in rat tracheal
epithelial cell cultures. In Vitro Cell Dev Biol Anim 1997; 33: 212–17.

29 Valcourt U, Kowanetz M, Niimi H, Heldin CH, Moustakas A.
TGF-beta and the Smad signaling pathway support transcriptomic
reprogramming during epithelial-mesenchymal cell transition.
Mol Biol Cell 2005; 16: 1987–2002.

30 Novotny L, Crha M, Rauser P, et al. Novel biodegradable
polydioxanone stents in a rabbit airway model.
J Thorac Cardiovasc Surg 2012; 143: 437–44.

31 Lischke R, Pozniak J, Vondrys D, Elliott MJ. Novel biodegradable
stents in the treatment of bronchial stenosis after lung
transplantation. Eur J Cardiothorac Surg 2011; 40: 619–24.

32 Vondrys D, Elliott MJ, McLaren CA, Noctor C, Roebuck DJ. First
experience with biodegradable airway stents in children.
Ann Thorac Surg 2011; 92: 1870–74.

  • Stem-cell-based, tissue engineered tracheal replacement in a child: a 2-year follow-up study
  • Introduction
    The recipient
    Pretransplant preparation
    Tracheal replacement surgery
    Role of the funding source


Biocompatibility of implants: lymphocyte/macrophage

James M. Anderson & Amy K. McNally

Received: 21 December 2010 /Accepted: 10 January 2011 /Published online: 27 January 2011
# Springer-Verlag 2011

Abstract The monocyte-derived macrophage is recognized
as a critical determinant in biocompatibility, but its
appearance in the chronic inflammatory phase is accompa-
nied by the presence of lymphocytes, which have been
much less studied in this regard. Here, we first present an
overview of the physiologic continuum comprising host
reactions to the surgical implantation of biomaterial.
Secondly, we describe our collective research efforts, which
indicate that lymphocytes are additional and key cellular
determinants of biocompatible outcome. Thus, bioengineer-
ing advances will require that lymphocyte responses be
regarded as integral components of innate inflammatory
and immune/immunotoxic cell interactions at sites of
biomaterial implantation.

Keywords Biomaterial . Foreign body response .

Inflammation . Lymphocyte . Macrophage

The inflammatory and wound healing responses
to biomaterials


The biocompatibility of implanted biomaterials is determined
by the degrees to which host homeostatic mechanisms are
perturbed during surgical placement of the implant and the

extents to which pathological consequences are created from
the ensuing inflammatory, wound healing, and foreign body
responses to surgical injury. The successful development of
novel biomaterials, tissue-engineered constructs, and prosthe-
ses will require in-depth mechanistic understanding of the
collective foreign body reaction.

Upon surgical implantation of a biomaterial, i.e., the
introduction of a foreign body into living host tissue, a
series of events ensues that is characterized by blood
protein adsorption, the arrival of acute and chronic
inflammatory cells, the formation of granulation tissue,
the foreign body reaction, and fibrosis or fibrous encapsu-
lation. This is outlined in Table 1. As the collective
responses of vascular tissue to injury, this continuum can
also be viewed as bimodal. The first phase involves the
activation of host mechanisms to limit or neutralize the
foreign body. The second phase is characterized by wound
healing and attempts to repair and reconstitute the injured
tissue at the implant site via the formation of fibroblastic
scar tissue, or combinations of both of these processes.

It must be emphasized that the driving force behind this
physiological continuum is the persistent presence of a non-
phagocytosable foreign body. However, each participating
cell type, and this is particularly known to be true for the
monocyte-derived macrophage, is differentially influenced by
chemically distinct material surfaces and by surface topogra-
phy. Therefore, to a great extent, it is biomaterial surface
chemistry that dictates the degrees to which events in the
continuum proceed, ultimately determining the overall bio-
compatible outcome of each biomaterial application.

Protein adsorption

Biomaterial surface property-dependent blood protein ad-
sorption occurs immediately upon surgical implantation of

This article is published as part of the Special Issue on Implanted
Devices: Biocompatibility, Tissue Engineering and Infection

J. M. Anderson (*) : A. K. McNally
Department of Pathology, Case Western Reserve University,
Wolstein Research Building Room 5104, 2103 Cornell Road,
Cleveland, OH 44106, USA

Semin Immunopathol (2011) 33:221–233
DOI 10.1007/s00281-011-0244-1

a biomedical device, and it is therefore this initial event that
precedes all subsequent sequelae. Thus, it is the blood
protein-modified biomaterial surface that all other host
components, including inflammatory and immune leuko-
cytes, encounter and/or interact with as an adhesion
substrate [1, 2]. The blood protein-modified biomaterial
surface is perhaps more appropriately referred to as a
provisional matrix, which, depending on biomaterial sur-
face chemistry, may also contain a milieu of chemokines,
cytokines, or other bioactive agents associated with
adsorbed blood protein components.

Although monocytes/macrophages express multiple
receptors with potential to mediate cell/substrate interac-
tions with blood proteins, we have identified both the β1
and β2 integrin families as necessary and sufficient
mediators of adhesion during monocyte-to-macrophage
development and foreign body giant cell (FBGC) formation
[3, 4]. Integrins comprise a large group of heterodimeric
transmembrane molecules that mediate both cell–extracellular
matrix and cell–cell interactions [5, 6]. These receptors are
well known as important mediators of adhesion signaling
between the extracellular and intracellular environments

In addition to their roles in leukocyte extravasation to
inflammatory sites, the β2 integrins αMβ2 and αXβ2 are
capable of interactions with multiple ligands to mediate
cell–particle or cell–substrate interactions and the induction
of β1 integrin expression [8, 9]. These include the blood
proteins complement C3bi and fibrin(ogen) adsorbed to
material surfaces [10–12]. Through the essential arginine–
glycine–aspartate cell attachment sequence, β1 integrins
can bind multiple protein ligands, including fibronectin,
vitronectin, collagens, and laminin; the interaction between
fibronectin and α5β1 integrin is a classic example [5].
Consistent with this, we have demonstrated that vitronectin-
adsorbed surfaces are an optimum substrate for monocyte
adhesion leading to macrophage development and FBGC
formation [13]. Therefore, the propensity for vitronectin
adsorption to biomaterials likely plays a critical role in the
development of the foreign body reaction.

Acute inflammation

The acute inflammatory response is marked by the presence
of short-lived blood-derived polymorphonuclear leukocytes
(PMNs), predominantly neutrophils, less than 10% eosino-
phils, and 1% basophils. Depending on the extent of injury
created by material implantation, PMNs that initially
rapidly accumulate in response to chemotactic stimuli cease
to emigrate to the site of injury within minutes to days.
PMNs are professional phagocytic cells that may function
to immediately remove foreign microorganisms [14].
Moreover, during this brief phase, their secretory activities
may be significant to the eventual biocompatible outcome.
In addition to the elaboration of reactive oxygen inter-
mediates, PMNs are capable of releasing cytokine media-
tors with potentials to influence the character and degree of
subsequent inflammatory cell recruitment and activation.
The resultant phenotypes of monocytes/macrophages and
wound healing cells in the chronic inflammatory response
to biomaterials may be, at least in part, governed by the
biomaterial-dependent behavior of PMNs. In this regard,
certain acute activities of PMNs have been studied on
several biomaterials [15–22]. We have also observed
that the co-culture of PMN together with monocytes/
macrophages diminishes the subsequent formation of
FBGC [23]. Related to this, the degranulation of tissue
mast cells is reported to mediate acute inflammatory
responses in vivo [24, 25]. However, the precise roles of
acute inflammatory leukocytes in influencing chronic
inflammatory, foreign body, and fibrosis host responses to
implanted materials remain largely unknown. Our inves-
tigations to date have demonstrated that multiple aspects of
the macrophage-mediated chronic response are influenced
by material surface chemistry. Therefore, we predict that
this may also be extended to the case of PMNs.

Chronic inflammation

Mononuclear leukocytes, i.e., monocyte-derived macro-
phages and lymphocytes, as well as the presence of plasma
cells, define the chronic response, which also features
neovascularization and the development of connective
tissue. With biocompatible materials, the chronic phase is
of limited duration and is usually evident for approximately
2 weeks. Macrophages are longer-lived than PMNs, and
they are also professional phagocytes with extraordinary
synthetic and secretory capacities [26]. Therefore, the
macrophage is believed to exert key controlling influences
on the ensuing wound healing/fibrosis responses. Depend-
ing on biomaterial surface chemistry, however, our studies
have indicated that adherent monocyte-derived macro-
phages are differently activated for cytokine production
and phenotypic expression. For example, we have identi-

Table 1 Biomaterial surface chemistry-dependent host reactions to
implanted biomaterials: sequence and continuum of events in the
inflammatory and wound healing responses

Surgical injury to host tissues/organs

Blood protein adsorption (provisional matrix formation)

Acute inflammation (PMNs)

Chronic inflammation (monocytes/macrophages, lymphocytes)

Granulation tissue (fibroblasts, vascular endothelial cells)

Foreign body reaction (macrophages, FBGCs, granulation tissue)

Fibrosis/fibrous encapsulation (development of fibrous capsule)

222 Semin Immunopathol (2011) 33:221–233

fied several key biomaterial-dependent chemokine and
cytokine mediators, interleukin (IL)-1β, IL-6, IL-8, and
tumor necrosis factor (TNF)-α, with potential to mediate
multiple autocrine and paracrine effects in the chronic
inflammatory and wound healing phases [27–30], as well as
a time-dependent switch in cytokine secretion from acute to
chronic phase phenotype [27].

The presence of lymphocytes in these scenarios is less
well studied, but our findings to date and those of others
suggest that lymphocytes play a far greater role in the host
response to implanted biomaterials than was previously
appreciated. This is more fully addressed below.

Plasma cells are a component of the humoral immune
system, and, although the role of these antibody-secreting
cells in biocompatibility is not known, it cannot be
completely discounted. Plasma cells have been reported to
be present at other sites of injury and inflammation [31],
indicating an as yet unknown significance in these

Granulation tissue

The initiation of wound healing is marked by the arrival of
fibroblasts and endothelial cells. These cell types are
recruited by the activities of chronic inflammatory macro-
phages, in the presence of which they proliferate to form
granulation tissue, so-called because of its characteristic
macroscopic granular appearance in healing wounds.
Granulation tissue develops only when there is tissue loss
or disturbance of tissue architecture such that the wound
healing process cannot fully restore the injured area to its
previous state. Thus, granulation tissue serves as the
intermediary substrate tissue for subsequent scar formation
or fibrosis. Within granulation tissue, endothelial cells
proliferate from pre-existing blood vessels and organize
into new capillaries in the process of angiogenesis to supply
the intermediary tissue with necessary blood flow. Fibro-
blasts synthesize extracellular matrix proteins, initially
proteoglycans and later collagens. Collagens form the basis
of the fibrous capsule (see below). Specialized fibroblasts
called myofibroblasts mediate wound contraction in gran-
ulation tissue.

Foreign body reaction

Another critical feature of the macrophage is its intriguing
and incompletely understood propensity to undergo fusion
with other macrophages, leading to the formation of
multinucleated giant cells. Giant cells are a classic
identifying feature of chronic inflammation arising from
the persistent presence of foreign microorganisms or
foreign bodies in multiple scenarios [32, 33]. Thus, the
multinucleated giant cells that form on or around biomate-

rials are referred to as foreign body giant cells or FBGC. In
the case of implanted materials, the presence of macro-
phages and/or FBGC together with the components of
granulation tissue is called the foreign body reaction (FBR).
The form and topography of the implanted material, e.g.,
the flat surface of breast prostheses versus the rough
topographies of vascular grafts, dictate the cellular compo-
sition of the FBR. In general, smooth or flat surfaced
biomaterials exhibit a thin layer of macrophages with
fibrosis compared to rough, high surface-to-volume fabrics,
particulates, and porous implants, which feature relatively
greater numbers of macrophages and FBGCs.

It is not uncommon for biomaterial-adherent FBGCs to
be quite large, with tens to even hundreds of nuclei and
measuring approximately 1 mm in diameter. These remark-
able giant cells have been observed to occupy up to 25% of
implant surface area on retrieved materials [34]. The FBR,
composed of macrophages and FBGCs, may persist for the
lifetime of the biomedical material or device [31, 33, 35].
Fibrosis (see below) also develops around the tissue/
material interface with its associated FBR. This fibrous
capsule appears to serve as a barrier between activities
related to the FBR and those of normal host tissue.

In the 1990s, biomaterial pitting and surface cracking
were discovered directly underneath adherent FBGCs in
vivo [36], and it was therefore inferred that FBGCs mediate
biomaterial degradation through the concentration of
phagocytic and oxidative activities at the giant cell/
biomaterial interface. Further, FBGC may produce cyto-
kines that bias wound healing cells toward a fibrogenic
phenotype. Consistent with this, we have also discovered
that cultures of fusing macrophages and FBGC strongly
express connective tissue growth factor (unpublished
results). Clinically, these activities appear to manifest as
biomaterial degradation, fibrous encapsulation, and device
failure [35, 37, 38].

In vitro, however, we originally discovered that FBGC
are induced from human monocyte-derived macrophages
by lymphokines that are well known to exert anti-
inflammatory effects on multiple monocyte/macrophage
activities [39, 40]. Later, we found that this is also the case
for α-tocopherol, the major component of the antioxidant
vitamin E [41]. This is more completely addressed below.
We further learned that lymphokine-induced FBGC forma-
tion is highly material surface property-dependent. For
example, IL-4-induced FBGC formation occurs readily on
nitrogenated and oxygenated but not on fluorinated or
silicone-coated model polystyrenes [42]. This phenomenon
has since been confirmed with multiple types of biomate-
rials and engineered surfaces [43–47].

With this perspective, FBGC formation has appeared to
be an undesirable phenomenon with a negative impact on
biocompatibility. It was speculated that FBGC formation

Semin Immunopathol (2011) 33:221–233 223

represented the joining of otherwise ineffective phagocytic
cell forces by host mechanisms attempting to remove the
implanted foreign body. Therefore, as presently perceived,
the foreign body reaction in terms of macrophage adhesion
and FBGC formation is believed to be undesirable and a
target for intervention. A broader perspective, however,
dictates that these phenomena represent partial aspects of a
continuum or collective host response that is initiated and
maintained by the introduction and persistent presence of a
non-phagocytosable foreign body. With this view, it is clear
that in order to fully understand the foreign body reaction,
additional host mechanisms responding to biomaterial
implantation must also be characterized.


The ideal biocompatible outcome for biomaterial implanta-
tion would clearly be the restoration of normal tissue
architecture and function following wound healing. With
only a few exceptions, however, fibrosis with fibrous
encapsulation is generally the final stage of the wound
healing response to biomaterials. In most cases, the
surgically injured tissue is composed primarily of perma-
nent cell types that cannot regenerate, and the process of
repair can provide only the fibrous capsule as replacement

Biomaterial-induced macrophage activation

Multinucleated giant cells, formed by the fusion of macro-
phages, are a hallmark feature of chronic inflammation due
to the persistent presence of a non-phagocytosable foreign
body. As mentioned above, FBGCs are believed to be
implicated in biomaterial degradation. However, of the
three macrophage fusion factors that we have identified (IL-
4, IL-13, and α-tocopherol) [39–41], each is well docu-
mented to down-modulate so-called “pro-inflammatory”
activities of macrophages. IL-4 and IL-13 are each
considered to be “alternative” macrophage activation
cytokines [48], and α-tocopherol exhibits host tissue
protective effects via antioxidant as well as non-
antioxidant mechanisms [49, 50]. Therefore, an apparent
paradox between biomaterial degradation and down-
modulation of inflammation by biomaterial-adherent mac-
rophages and FBGCs remains to be resolved, i.e., what is
the biomaterial-induced macrophage activation phenotype?

The concept of “alternative” macrophage activation was
introduced to distinguish Th2 lymphokine (IL-4)-activated
macrophages from “classically” activated macrophages,
which develop under the influence of the Th1 lymphokine
interferon (IFN)-γ [51–54]. IL-4 and IFN-γ are well
documented to exert antagonistic effects on macrophages.
For example, mannose receptor-mediated phagocytosis is

strongly induced by IL-4 and inhibited by IFN-γ, which
instead supports IgG-mediated phagocytosis. The produc-
tion of pro-inflammatory cytokines (IL-1, IL-6, TNF-α) is
induced by IFN-γ but inhibited by IL-4, which instead
induces anti-inflammatory cytokines (IL-1 receptor antag-
onist and IL-10), thereby promoting wound healing and
matrix deposition. Microbicidal reactive oxygen and nitro-
gen species are induced by IFN-γ, but IL-4 induces
arginase-1 activity, which mediates collagen deposition
and tissue repair. Further, our investigations have revealed
that IL-4, IL-13, or α-tocopherol promotes the formation of
very large foreign body-type multinucleated giant cells that
are morphologically identical to those found on biomate-
rials [39–41]. This is illustrated in Fig. 1. For purposes of
contrast, Fig. 1 also shows that, although IFN-γ together
with IL-3 or granulocyte-macrophage colony-stimulating
factor also induces macrophage fusion, the resulting
multinucleated giant cells are the morphologically distinct
Langhans-type giant cells [39], which are not found
adherent to retrieved biomaterials. The functional signifi-
cances of these morphological variants at various sites of
chronic inflammation have yet to be established.

Our discoveries that the alternative macrophage activa-
tion cytokines IL-4 and IL-13 are potent inducers of
macrophage fusion in vitro [39, 40] led to the finding that
IL-4 plays an important role in FBGC formation in vivo
[55]. IL-4 and IL-13 are well-recognized as Th2 lympho-
cyte products. However, IL-4 and IL-13 are also produced
by natural killer (NK) and NKT lymphocytes. Alternatively,
eosinophils, basophils, and mast cells are also potential

Fig. 1 Alternative versus classical macrophage activation leads to
morphological variants of multinucleated giant cells in vitro. The
alternative activation cytokine IL-4 (or IL-13 or α-tocopherol)
promotes macrophage fusion and the formation of large foreign
body-type giant cells (FBGC) with randomly arranged nuclei and high
degrees of cytoplasmic spreading. The classical activation cytokine
IFN-γ (together with IL-3 or granulocyte-macrophage colony-
stimulating factor) induces more limited degrees of macrophage
fusion with resultant Langhans-type giant cells [39–41]

224 Semin Immunopathol (2011) 33:221–233

sources of these cytokines [56]. The precise origins of
FBGC-inducing cytokines at the implant site remain

More recently, the repertoire of macrophage activation
has been expanded to include a state of “innate” activation
in response to Neisseria, which exhibits increased produc-
tion of reactive oxygen and nitrogen species, phagocytosis,
and expression of co-stimulatory molecules [53]. Therefore,
as studies on macrophage activation advance, it is likely
that macrophage heterogeneity will be analogously ad-
vanced to reveal multiple phenotypic states of macrophage
activation customized to distinct stimuli, including implanted

Our more recent in vitro studies have revealed several
novel characteristics of IL-4-induced fusing macrophages/
FBGCs. Namely, they strongly express an alternative
activation marker, galectin-3, the galectin-3 ligand and
lymphocyte-co-stimulatory molecule CD98, as well as
additional and select lymphocyte-co-stimulatory molecules
that may indicate down-modulation of lymphocyte activi-
ties (unpublished findings). These findings will be extended
toward a more comprehensive characterization of the
macrophage activation phenotype that ensues from inter-
actions with implanted biomaterials, biomedical devices,
and prostheses. Thereby, we may ultimately illuminate the
physiological significance of FBGC formation at sites of
biomaterial implantation. This, in turn, will provide
clinicians with a more complete perspective on the efficacy
of interventions in the foreign body reaction at implant

Inflammatory cell interactions in biocompatibility

Macrophages as key determinants of biocompatible

Inflammatory cell interactions significantly impact the
biocompatibility and function of medical devices, prosthe-
ses, and biomaterials and can ultimately precipitate their
clinical failure [1]. As subjects of extensive investigations,
monocyte-derived macrophages and FBGCs are believed to
exert critical effects on both tissue and implanted bioma-
terial. Adherent macrophages and/or FBGCs may promote
clinically significant biomaterial biodegradation through the
release of reactive oxygen intermediates with resultant
oxidative cleavage of the soft segment in polyether polyur-
ethanes. A classic example is the apparent macrophage and/
or FBGC-mediated biodegradation of Pellethane 80A used
as insulation in pacemaker leads [38]. Over time, this
material undergoes oxidative biodegradation with loss of
insulation integrity and ultimate pacemaker failure. A
second example is the focal biodegradation of Dacron®

polyester vascular grafts. Biodegradation occurs over long
time periods with areas of material loss being replaced by
fibrous tissue. The significant clinical outcome is focal
aneurysm in the vascular graft wall, necessitating replace-
ment surgery with its attendant problems [35].

Adherent macrophages and FBGC also possess poten-
tials to impact host tissue responses via the production of
multiple chemokines and cytokines that promote cellular
accumulation and proliferation. Clinically, these phenome-
na are well known and documented to lead to failure
mechanisms. Examples include anastomotic hyperplasia in
human vascular grafts leading to blood flow turbulence and
thrombosis and coronary artery stent restenosis, in which
the inflammatory response to both simple and polymer-
coated drug eluted stents can result into smooth muscle cell
and fibroblast proliferation in the stenotic component [57,
58]. An additional example is found in aseptic loosening of
joint prostheses, as in total hip or knee replacement, where
the fibrotic response separates bone from either the
prosthesis or its polymethylmethacrylate bone cement
[35]. Fibrous capsule formation in this area then leads to
micromotion that, in turn, causes aseptic loosening of the
prosthesis. Related to this, one mechanism for breast
implant failure involves capsular contraction, leading to
pain and breast deformation [37]. Biosensor function is also
compromised by biofouling and fibrous capsule formation
[59]. Thus, fibroblast proliferation and enhanced extracel-
lular matrix formation have been implicated clinically in the
failure of a wide variety of medical devices and prostheses.

Modulation of adhesion and activation of macrophages
and FBGC with anti-inflammatory drugs may enhance the
performance of pacemaker lead electrodes by reducing the
inflammatory response and ultimate fibrous encapsulation.
However, in clinical application, this type of solution to
reduce local inflammatory cell interactions is rare and
limited. Long-term systemic use of anti-inflammatory drugs
may compromise host defenses, leading to impaired
inflammatory and immune responses and increased potential
for infection.

Synthetic and modified natural tissue engineering scaf-
folds are also susceptible to the foreign body reaction,
which can markedly reduce the function and usefulness of
tissue-engineered constructs [60]. Predictably, surface mod-
ifications that inhibit, attenuate, or alter the progression of
these events will enhance the function of these constructs.
We believe that material-dependent monocyte/macrophage
adhesion mechanisms, adhesion signaling, anoikis (apopto-
sis induced by adhesion failure), and chemokine/cytokine
networks are four key acute and chronic cellular response
mechanisms that, once understood, may potentially be
targeted for control of biocompatible outcome. It is widely
believed that attenuation of these critical monocyte,
macrophage, or FBGC responses will lead to a diminished

Semin Immunopathol (2011) 33:221–233 225

foreign body reaction and increased efficacy of medical
device function.

The emerging roles of lymphocytes in biocompatibility

It is important to re-emphasize that lymphocytes appear
together with monocyte-derived macrophages in the
chronic phase. Therefore, their potential interactions
with and influences on macrophages as well as the
effects of biomaterials on lymphocytes themselves must
be considered in our efforts to understand biocompati-
bility mechanisms.

Lymphocyte responses to conventional microbial patho-
gens have long been extensively studied within the context
of classical immunity, which exhibits specificity and
memory. Biomaterials, largely synthetic polymers or met-
als, are apparently not immunogenic in the classic adaptive
sense. Thus, an awareness that lymphocytes react to a
foreign body outside of this classical context is emerging.
In this regard, we can utilize known aspects of lymphocyte
biology to investigate their responses to biomaterials, but
we must broaden developing perceptions on biocompati-
bility mechanisms to encompass possibilities beyond
classical immune responses. Our recent studies, as dis-
cussed below, have demonstrated that lymphocytes play a
much greater role in the inflammatory and foreign body
responses to implanted biomaterials than was previously
realized [61–67].

Also in support of this, lymphocyte responses to
synthetic materials have been clinically observed. Patients
implanted with left ventricular assist device (LVAD)
experience a reduction in circulating CD4+ T lymphocyte
populations, which is due to their apoptosis [68]. These
patients also have a B lymphocyte hypersensitivity, which
complicates their ability to receive a donor heart and has
been found to be due to T lymphocyte activation of B cells.
The culture of human peripheral mononuclear cells from
healthy donors with LVAD biomaterial is reported to result
in T lymphocyte activation. Intriguingly, activated T
lymphocytes adherent to macrophages are present on LVAD
surfaces retrieved from patients [69].

NK lymphocytes, which play prominent roles in host
defense against viral infection, have also been found to be
adversely affected by exposure to biomedical polymers
materials. Patients undergoing hemodialysis exhibited
reduced numbers of circulating NK cells after dialysis.
Moreover, this finding was extended to demonstrate
subsequent decreased NK cell activity in vitro [70, 71].
Related to this, hemodialysis patients acquire an increased
susceptibility to and incidence of viral infections [72].
These collective findings clearly indicate that lymphocyte
responses to biomedical polymers can lead to adverse
clinical consequences and reduced biocompatibility.

Conventionally, lymphocytes become activated through
interactions with antigen-presenting cells, i.e., macrophages
and dendritic cells, which present processed antigen bound
to major histocompatibility complex molecules on their cell
surfaces [73, 74]. Characteristics of activation include
expression of cell surface markers of activation, calcium
flux, and production of the classic activation lymphokines
IL-2 and IFN-γ. The activation markers CD40 ligand and
calcium influx are induced on T lymphocytes cultured with
polyurethane [75]. Activation alone, however, is not
sufficient for effector function. T lymphocytes require
additional cognate interactions with specific molecules on
antigen-presenting cells, e.g., macrophages; these are
collectively referred to as co-stimulation. Following cog-
nate activation, lymphocytes may undergo clonal expansion
by proliferation [73]. Specific types of co-stimulatory
molecules also regulate functional outcomes of lymphocyte
activation; lymphocyte effector function may be up- or
down-regulated [76]. Alternative to cognate activation, in
the presence of activation signals without co-stimulation,
lymphocytes cannot acquire effector function and are
referred to as anergic. Anergic responses are a mechanism
for suppression of inappropriate immune reactivity, such as
to self antigen [77]. A further alternative to activation is
apoptosis. By this mechanism, cells which may have been
inappropriately activated undergo programmed cell death
and are removed by macrophages [78]. The clinical
findings outlined above indicate that mechanisms of
lymphocyte activation or anergy, as well as apoptosis are
precipitated by lymphocyte interactions with implanted

Macrophage/lymphocyte interactions

The progression of inflammatory events and the foreign
body reaction features the appearance of mononuclear
leukocytes (monocytes/macrophages and lymphocytes) at
the site of biomaterial implantation. The guided movements
of these cells are mediated by cytokines and chemokines,
which also participate in lymphocyte differentiation, lym-
phocyte proliferation, and macrophage activation. Media-
tors such as I-309, macrophage inflammatory protein
(MIP)-1β, IL-8, MIP-3α, and IFN-γ-induced protein (IP)-
10 can attract specific lymphocyte subpopulations [79] or
activate/enhance activation of lymphocytes (IL-1β, MIP-1β,
IL-6) [80, 81].

Lymphocytes and monocytes/macrophages are capable
of activating each other through direct and indirect
mechanisms. T lymphocytes can be activated by cytokines
such as IFN-γ, IL-2, TNF-α, IL-6, IL-15, and IL-18 [82].
NK lymphocytes are activated and proliferate when
exposed to chemokines MIP-1β, macrophage chemotactic
protein (MCP)-1, and chemokine C-C motif ligand (CCL) 5

226 Semin Immunopathol (2011) 33:221–233

[81]. Activated T lymphocytes induce production of the
pro-inflammatory cytokines IL-1, TNF-α, and IL-6, and
the chemokines IL-8, MCP-1, and MIP-1β from mono-
cytes/macrophages in a contact-dependent manner [83,
84]. Depending on the cytokine/chemokine stimulus,
macrophages can be activated to produce a myriad of
cytokines [85].

Cytokines that we have detected at sites of biomaterial
implantation are also capable of suppressing lymphocyte
and macrophage behaviors. IL-10 is an anti-inflammatory
cytokine capable of inducing CD4+ T cell antigen-specific
anergy [86]. Chronic TNF-α exposure can lead to T cell
hypo-responsiveness by down-regulation of the T cell
receptor/CD3 [87]. TNF-α also induces apoptosis [88].
We have demonstrated induction of macrophage apoptosis
on biomaterial surfaces through TNF-α-mediated mecha-
nisms and inhibition of these mechanisms by the lympho-
kine IL-4 [89, 90]. This indicates further influences of
lymphocytes in the complex cell–cell interactions on
biomaterials and presents an additional avenue for investi-
gation of the mechanisms by which lymphocytes influence
critical behaviors of macrophages and FBGCs.

NK and NKT lymphocytes are key mediators of innate
host response mechanisms through cytokine production and
regulation of autoimmune and inflammatory diseases [91–
94]. NK cells are activated by direct receptor/ligand
interactions but also indirectly by cytokines, for example,
to produce IFN-γ [93, 95]. As mentioned above, there is
evidence that NK lymphocyte functions are affected by
exposure to dialysis membranes [70–72]. However, their
responses to implanted biomaterial surfaces have not been
characterized and remain unknown. Importantly, these cell
types are potential sources of IL-4 and/or IL-13 [92, 96–
98], which may promote the foreign body reaction by
inducing macrophage fusion to form FBGCs [39, 40].

Our current data suggest that T lymphocyte-deficient
mice exhibit a normal foreign body reaction to implanted
material (see below) [99]. Therefore, whether NK or NKT
lymphocytes are sources of IL-4/IL-13 relevant to FBGC
formation remains an intriguing question. In addition, IL-13
produced by NKT lymphocytes has been implicated in the
progression of fibrosis [100, 101]. Therefore, the potential

participation of this lymphocyte subset in the fibrous
encapsulation of implanted materials, perhaps via IL-13
stimulation of macrophages to produce transforming
growth factor (TGF)-β [102], is another critical issue.

Our investigations to date have revealed further evidence
that, in addition to the macrophage, it is important to consider
the influences of lymphocytes on biocompatible outcome. We
have discovered lymphocyte effects on macrophage adhesion
and fusion as well as on biomaterial-dependent cytokine
production. Interestingly, the secondary response of lympho-
cytes to biomaterials has been found to be distinct from the
primary response in vivo. The significant findings from our
investigations on lymphocyte/macrophage interactions are
summarized in Table 2 and further discussed below.

Lymphocyte modulation of macrophage adhesion
and FBGC formation

Despite their early and relatively transient presence at sites
of biomaterial implantation, few studies have focused on
how lymphocytes may influence the subsequent foreign
body response. Based on our in vitro human monocyte/
macrophage/FBGC system, a lymphocyte/macrophage co-
culture system was developed to address these unknowns
[61]. We discovered that when lymphocytes are present at
an optimal ratio of 25:1 during the initial adhesion of
monocytes, the rates of monocyte adhesion and subsequent
macrophage fusion are significantly increased (50–60%
fusion) when compared to monocytes alone (10% fusion).
If lymphocytes are added at later time points (days 3 or 7),
these differences are diminished, indicating that lympho-
cyte interactions with monocytes during early stages of
adhesion and culture are critical for increasing monocyte/
macrophage adhesion and subsequent macrophage fusion.
Importantly, we found that 90% of adherent lymphocytes
associate with adherent macrophages and not with bioma-
terial surfaces. In turn, these interactions led to increases in
lymphocyte proliferation, which was greatest on day 3 but
also significantly greater than a lectin-stimulated lympho-
cyte control population at later time points (days 7 and 10).
These studies are the first to demonstrate that interactions
between monocytes/macrophages and lymphocytes pro-

Promote macrophage adhesion and fusion to form FBGCs

Lymphocyte/macrophage interactions promote lymphocyte proliferation

Lymphocyte/macrophage interactions are influenced by biomaterial chemistry

Lymphocyte/macrophage interactions promote biomaterial-dependent cytokine production

Lymphocyte/macrophage interactions promote interferon-γ production

Lymphocyte secondary responses to biomaterials recruit T lymphocytes and phagocytes

In vivo cytokine profiles are biomaterial-dependent

Nude (Th1/Th2-deficient mice) demonstrate a normal foreign body reaction

Table 2 The roles of
lymphocytes in biocompatibility

Semin Immunopathol (2011) 33:221–233 227

voke responses from each cell type that are directly relevant
to the foreign body reaction.

Lymphocyte/macrophage interactions are differentially
influenced by material surface chemistry

Using our lymphocyte/macrophage co-culture system, we
addressed the effects of different material surface
chemistries on monocyte adhesion, macrophage fusion,
and lymphocyte proliferation [62]. A series of poly
(ethylene terephthalate) (PET)-based materials was
employed to present hydrophobic, hydrophilic/neutral,
hydrophilic/anionic, and hydrophilic/cationic chemistries
to cells. Hydrophilic/neutral surfaces interfered with initial
monocyte/macrophage adhesion (day 0), thereby preclud-
ing subsequent macrophage fusion and FBGC formation.
These surfaces also were unable to support lymphocyte
proliferation at later time points. In contrast, hydrophilic/
anionic surfaces exhibited decreases in macrophage adhe-
sion between days 3 and 7, but increases lymphocyte
proliferation and macrophage fusion at later time points.
Hydrophilic/cationic surfaces also demonstrated decreases
in macrophage adhesion between days 3 and 7, but, unlike
anionic materials, there were no corresponding increases
in lymphocyte proliferation or macrophage fusion on the
cationic surfaces. These results clearly demonstrate that
differences in material surface chemistry are capable of
provoking very different cellular responses. Effects on
lymphocyte function as well as monocyte/macrophage
adhesion and fusion resulting from interactions with
lymphocytes are readily apparent in this model system.

Chemokines/cytokines in lymphocyte/macrophage

The progression/continuum of inflammatory events and the
foreign body reaction (Table 1) feature the appearance of
mononuclear leukocytes (monocytes/macrophages and lym-
phocytes) at the site of biomaterial implantation. The
guided movements of these cells are mediated by cytokines
and chemokines, which also participate in lymphocyte
differentiation, lymphocyte proliferation, and macrophage
activation [54, 80, 103]. Our studies point to cytokine-
mediated (indirect) signaling in lymphocyte enhancement
of monocyte adhesion and macrophage fusion [61].

Utilizing cytokine protein arrays and ELISA to investi-
gate cytokine and chemokine production from lymphocyte/
macrophage co-cultures, we identified production of solu-
ble mediators that are capable of targeting both lympho-
cytes and macrophages [63]. For instance, MCP-1, a
chemoattractant for monocytes/macrophages, also supports
macrophage fusion [104]. Other detected mediators (I-309,
MIP-1β, IL-8, MIP-3α, and IP-10) can attract specific

lymphocyte subpopulations [79] or activate or enhance
activation of lymphocytes (IL-1β, MIP-1β, IL-6) [80, 81].

Lymphocytes and monocytes/macrophages are capable
of activating each other through direct and indirect
mechanisms. T lymphocytes can be activated by cytokines
such as IL-2, TNF-α, IL-6, IL-15, and IL-18 [82]. NK cells
are activated and proliferate when exposed to the chemo-
kines MIP-1β, MCP-1, and CCL5 [81]. Activated T
lymphocytes induce production of the pro-inflammatory
cytokines IL-1, TNF-α, and IL-6, and the chemokines
IL-8, MCP-1, and MIP-1β from monocytes/macrophages
in a contact-dependent manner [83, 84]. Depending on the
cytokine/chemokine stimulus, macrophages can be acti-
vated to produce multiple cytokines [85]. The specific
soluble mediator(s), lymphocyte/macrophage cell surface
molecular interactions, and lymphocyte subpopulation(s)
involved in the observed enhancement of adherent
monocyte, macrophage, or FBGC formation have been
targets in our investigations.

Cytokines that we have detected at sites of biomaterial
implantation are also capable of suppressing lymphocyte
and macrophage behaviors. IL-10 is an anti-inflammatory
cytokine capable of inducing CD4+ T cell antigen-specific
anergy [86]. As mentioned earlier, chronic stimulation by
TNF-α down-modulates the T cell receptor/CD3 and
induces T cell hypo-responsiveness [87]. TNF-α also
induces apoptosis [88]. Our studies have shown that the
biomaterial-dependent induction of macrophage apoptosis
occurs through TNF-α-mediated mechanisms; these mech-
anisms are blocked by IL-4 [46, 89, 90]. The significance
of these cytokines in modulating lymphocyte behavior at
sites of biomaterial implantation is not yet clear.

Lymphocyte/macrophage interactions
and biomaterial-dependent cytokine production

To investigate the effects of biomaterial surface chemistry
on the production of cytokines, chemokines and extracel-
lular matrix metalloproteinases (MMPs) from lymphocytes
and macrophages, human monocytes and lymphocytes were
co-cultured on chemically distinct PET-based material
surfaces as described above. Antibody array screening
indicated that the majority of detected proteins are
inflammatory mediators that guide the early inflammatory
phases of wound healing [63]. Proteomic ELISA quantifi-
cation and adherent cell analysis were performed after 3, 7,
and 10 days of culture. IL-2 was not detected in any co-
cultures suggesting that lymphocyte activation does not
occur by classic immune response mechanisms. The
hydrophilic/neutral surfaces increased IL-8 relative to the
hydrophobic PET surface (p<0.05). The hydrophilic/anionic surfaces promoted increased TNF-α over hydrophobic and cationic surfaces and increased MIP-1β compared to

228 Semin Immunopathol (2011) 33:221–233

hydrophobic surfaces (p<0.05). Since enhanced macrophage fusion was observed on hydrophilic/anionic surfaces, the production of these cytokines may be related to FBGC formation. The hydrophilic/cationic surface promoted IL-10 production and increased MMP-9/tissue inhibitor of MMP (TIMP) relative to hydrophilic/neutral and anionic surfaces (p<0.05). The collective results of this study suggest that hydrophilic/neutral and anionic surfaces promote pro- inflammatory responses and reduced matrix degradation, whereas the hydrophilic/cationic surfaces induce an anti- inflammatory response and greater MMP-9/TIMP with an enhanced potential for matrix breakdown [63]. This in vitro investigation also underscores the usefulness of protein arrays in assessing the roles of soluble mediators in the inflammatory response to biomaterials and may provide perspective for future clinical assessments. How specific lymphocyte and macrophage interactions result in biomaterial-dependent cytokine production is unclear.

Direct versus indirect lymphocyte/macrophage interactions
and biomaterial-dependent production of selected cytokines

Lymphocyte interactions with adherent macrophages and
model surface-modified materials were further investigated
by culturing monocytes alone or together with lympho-
cytes, either in direct co-cultures or indirectly in transwells
with 0.02-μm pores to allow diffusion of soluble compo-
nents [64]. The cultures were carried out on PET-based
photograft co-polymerized material surfaces displaying
distinct hydrophobic, hydrophilic/neutral, hydrophilic/
anionic, and hydrophilic/cationic chemistries, with mono-
cytes/macrophages seeded onto the test materials in the
transwell system. After periods of 3, 7, and 10 days,
cytokine production was quantified by ELISA and normal-
ized to adherent macrophage/FBGC density to yield a
measure of adherent macrophage/FBGC activation. Hydro-
philic/neutral and hydrophilic/anionic surfaces evoked the
highest levels of activation. Adherent macrophages/FBGCs
in co-culture with lymphocytes increased production of IL-
1β, TNF-α, IL-6, IL-8, and MIP-1β on the base PET,
hydrophobic, and hydrophilic/cationic surfaces indicating a
role for lymphocytes in the inflammatory response to
biomaterials. IL-10 production was material-dependent but
unaffected by direct or indirect lymphocyte interactions. At
early time points (day 3), indirect signaling promoted
enhanced macrophage/FBGC activation while direct inter-
actions may inhibit that response. At later time points
(10 days), direct interactions dominated and increased the
activation level. Lymphocytes did not have a significant
effect on MMP-9, TIMP-1, and TIMP-2 production.
Biomaterial surface chemistries differentially affected lym-
phocyte and macrophage/FBGC interactions as the in-
creased levels of activation were not evident on

hydrophilic/neutral or entirely evident on hydrophilic/
anionic surfaces. Therefore, although the majority of
adherent lymphocytes in co-cultures are in direct contact
with adherent macrophages and not with materials
lymphocyte-mediated macrophage activation for cytokine
production does not require direct interactions with lym-
phocytes. Whether direct interactions between these cell
types actually inhibit the macrophage response is an
intriguing question.

Lymphocyte/macrophage interactions induce
the lymphokine interferon (IFN)-γ

Interestingly, our most recent data reveal that IFN-γ is
produced in significant picogram per millilter amounts in
lymphocyte/macrophage co-cultures on days 3, 7, and 10,
either in direct co-cultures or when these populations are
segregated in transwell cultures [65]. Essentially, no IFN-γ
is detected in lymphocyte-only (or monocyte-only) cultures,
suggesting that either direct or indirect interactions of
lymphocytes with biomaterial-adherent monocyte-derived
macrophages are required for the induction of this classical
activation lymphokine. We interpret this to mean that IFN-
γ production from lymphocytes requires induction by a
macrophage-derived cytokine(s) or other soluble factor, but
we also observed that the levels of IFN-γ in direct co-
cultures is increased relative to those in indirect transwell
cultures, indicating synergism of direct and indirect
interactions. In addition, IFN-γ production is model surface
material-dependent as follows: hydrophilic/anionic≥
hydrophobic=base PET≥hydrophilic/cationic≥hydrophilic/
neutral. Our earlier work demonstrated that IFN-γ in
combination with macrophage developmental cytokines
could induce limited macrophage fusion and multinucleated
giant cell formation [39]. Now, we have discovered that
lymphocyte interactions with biomaterial-adherent macro-
phages elicit the production of this classic macrophage
activation cytokine from lymphocytes in vitro. As noted
above, IL-2 was not detected in lymphocyte/macrophage
co-cultures, indicating selective lymphokine signaling via
IFN-γ. These data provide further evidence for important
interactions between lymphocytes and macrophages at sites
of biomaterial implantation.

In vivo secondary responses to biomaterial implantation
are characterized by recruitment of T lymphocytes
and phagocytic cells

Biomaterials are widely believed to be non-immunogenic,
as polymer subunits are not considered to be antigens in the
classic sense. Therefore, the nature of host response to
implanted polymeric biomedical materials has been largely
assumed to be non-adaptive or non-immune. Classic

Semin Immunopathol (2011) 33:221–233 229

features of adaptive immunity are specificity and memory,
i.e., specific antigen-driven lymphocyte responses which
are more rapid and magnified upon secondary exposure,
with increases in the percentages of CD4+ (helper) and
CD4+/CD25+ (activated) T lymphocytes. To directly
address this issue, we compared the in vivo primary and
secondary host responses to three clinically-relevant bio-
materials: a poly (ether urethane) (PEU), silicone rubber,
and PET using our rat cage implant model [66]. Caged
polymer samples were subcutaneously implanted for
14 days and then explanted. After a 2-week healing period,
the identical rats were implanted a second time with cages
containing new samples of the same polymers for an
additional 2 weeks. Exudates were analyzed at 4, 7, and
14 days post-primary or post-secondary implantation by
flow cytometry for the following cell types: CD4+ T
lymphocytes, CD4+/CD25+ T lymphocytes, CD8+ T
lymphocytes (suppressor), B lymphocytes, PMNs, and
macrophages. At day 14 following secondary implantation,
we observed significant increases in T lymphocytes, PMNs,
and macrophages in the exudates compared to primary
implantation for all groups. B lymphocytes were not
detected. Significantly, CD4+/CD8+ ratios, percentages of
CD4+/CD25+ T lymphocytes, macrophage adhesion, and
FBGC formation were each comparable between primary
and secondary implantations. Therefore, our results argue
against an adaptive host response in the classic sense.
However, following re-implantation of biomaterials, we
found significant quantitative increases in both phagocytic
cells and T lymphocytes with constant T cell subset
distributions. This indicates recruitment of these cell types
upon secondary exposure to biomedical polymers. Whether
the observed secondary cellular recruitment is biomaterial-
specific remains to be determined

In vivo cytokine profiles are biomaterial-dependent

To identify cytokine signals in cell–cell communication
during the foreign body reaction in vivo, investigation of
cytokines at biomaterial implant sites was carried out [67].
Macrophage activation cytokines (IL-1β, IL-6, TNF-α),
cytokines important for macrophage fusion (IL-4, IL-13), a
T cell activation cytokine (IL-2), anti-inflammatory cyto-
kines (IL-10, TGF-β), and chemokines (GRO/KC, MCP-1)
were quantified at biomaterial implant sites using a
multiplex immunoassay and ELISA. Empty cages (con-
trols) or cages containing synthetic biomedical polymer,
either PEU, silicone rubber, or PET, were implanted
subcutaneously in rats for 4, 7, or 14 days, and cytokines
in exudate supernatants and macrophage surface adhesion
and fusion were quantified. The presence of a polymer
implant did not affect the levels of IL-1β, TGF-β, and
MCP-1 in comparison to the control group. IL-2 was not

detected in these exudate samples. However, the levels of
IL-6, TNF-α, IL-4, IL-13, IL-10, and GRO/KC were
modulated by polymer implantation. The levels of IL-6
and TNF-α were significantly greater in animals implanted
with PEU and silicone rubber, which are materials that,
conversely, do not support FBGC formation. These results
confirm and extend our in vitro data to indicate that the
cytokine production profiles of adherent macrophages and
FBGCs in vivo are biomaterial-dependent. Further, the
possibility is raised that the presence of macrophage fusion
and FBGC formation on biomaterial surfaces represents
host down-modulation of pro-inflammatory cytokine pro-
duction, perhaps via phagocytic removal/sequestration of
macrophages actively elaborating these cytokines. If so,
current perception of the so-called “inflammatory giant
cell” must evolve to accommodate this new concept.

Nude mice exhibit a normal foreign body reaction
to implanted material

Our data with nude (i.e., athymic) mice indicate that these
Th1 and Th2 lymphocyte-deficient mice are capable of
forming FBGCs that are comparable to those of normal
mice both in morphology and extent of formation [99].
These data suggest that Th2 lymphocytes are not the source
or not the only source of relevant fusion-inducing IL-4 or
IL-13. This is consistent with the common theme of
redundancy in host defense systems. Other potential
sources of IL-4 and IL-13 are NK or invariant NKT
lymphocytes (which do not undergo maturation in the
thymus), mast cells, basophils, and eosinophils [56, 98].
Furthermore, although our results confirm that IL-4 plays
an important role in FBGC formation on biomaterials in
vivo [55], it is very possible that additional macrophage
fusion-inducing cytokines will be discovered. For example,
IL-21 is a more recently discovered cytokine which exerts
effects that are similar to IL-4 and IL-13 in the alternative
activation of macrophages [105]. Whether IL-21 can induce
FBGC formation is not yet known. Therefore, the questions
raised by these preliminary results are intriguing and
significant. Their answers will definitively expand our
current understanding of the inflammatory response to
implanted materials.


Host response evaluation of implants is critical to deter-
mining the safety and biocompatibility of medical devices,
prostheses, and biomaterials. This chapter first presents an
overview of the sequence and continuum of events in the
inflammatory and wound healing responses that facilitate
biocompatibility evaluation. This is followed by summaries

230 Semin Immunopathol (2011) 33:221–233

of our results to date on lymphocyte responses to
biomaterials in vitro and in vivo. Our investigations
indicate that lymphocytes and lymphocyte/macrophage
interactions play much more significant roles at implant
sites than have been previously appreciated. Given the
current and future development of tissue-engineered con-
structs and bioactive agent delivery systems, lymphocyte
function and immunotoxicity evaluation may play a
significant role in determining the safety of these respective
systems. This, in turn, requires a better understanding of the
mechanisms leading to immunotoxicity or acquired immu-
nity as they relate to the specific composition of the
respective tissue-engineered constructs and/or bioactive
agent delivery systems under investigation. An understand-
ing of the mechanisms and methods presented in this
chapter can permit the early identification of factors that
may compromise or obviate the biocompatibility of medical
devices, prostheses, and biomaterials. Early identification
of problems related to biocompatibility including immuno-
toxicity can permit new design criteria to be introduced into
the research development process.

Acknowledgment The authors gratefully acknowledge the support
of the National Institutes of Health, Institute of Biomedical Imaging
and Bioengineering, Grant EB-00282.


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DOI: 10.4068/cmj.2011.47.1.1
Ⓒ Chonnam Medical Journal, 2011 Chonnam Med J 2011;47:1-131

Review Article

Tissue Engineering: Current Strategies and Future Directions
Jennifer L. Olson, Anthony Atala and James J. Yoo*

Wake Forest Institute for Regenerative Medicine, Wake Forest University School of Medicine, NC, USA

Novel therapies resulting from regenerative medicine and tissue engineering technol-
ogy may offer new hope for patients with injuries, end-stage organ failure, or other clin-
ical issues. Currently, patients with diseased and injured organs are often treated with
transplanted organs. However, there is a shortage of donor organs that is worsening
yearly as the population ages and as the number of new cases of organ failure increases.
Scientists in the field of regenerative medicine and tissue engineering are now applying
the principles of cell transplantation, material science, and bioengineering to construct
biological substitutes that can restore and maintain normal function in diseased and
injured tissues. In addition, the stem cell field is a rapidly advancing part of re-
generative medicine, and new discoveries in this field create new options for this type
of therapy. For example, new types of stem cells, such as amniotic fluid and placental
stem cells that can circumvent the ethical issues associated with embryonic stem cells,
have been discovered. The process of therapeutic cloning and the creation of induced
pluripotent cells provide still other potential sources of stem cells for cell-based tissue
engineering applications. Although stem cells are still in the research phase, some
therapies arising from tissue engineering endeavors that make use of autologous, adult
cells have already entered the clinical setting, indicating that regenerative medicine
holds much promise for the future.

Key Words: Biomaterials; Cell transplantation; Regenerative medicine; Stem cell; Tissue

This is an Open Access article distributed under the terms of the Creative Commons Attribution Non-Commercial
License ( which permits unrestricted non-commercial use,
distribution, and reproduction in any medium, provided the original work is properly cited.

Article History:
received 31 March, 2011
accepted 8 April, 2011

Corresponding Author:
James J. Yoo
Wake Forest Institute for Regenerative
Medicine, Wake Forest University
School of Medicine, Medical Center
Boulevard, 1834, Wake Forest Road,
Winston-Salem, NC 27109, USA
TEL: +1-336-713-7294
FAX: +1-336-713-7290


 Patients suffering from diseased and injured organs are
often treated with transplanted organs, and this treatment
has been in use for over 50 years. In 1955, the kidney be-
came the first entire organ to be replaced in a human, when
Murray transplanted this organ between identical twins.
Several years later, Murray performed an allogeneic kid-
ney transplant from a non-genetically identical patient in-
to another. This transplant, which overcame the immuno-
logic barrier, marked a new era in medicine and opened the
door for use of transplantation as a means of therapy for
different organ systems.
 As modern medicine increases the human lifespan, the
aging population grows, and the need for donor organs
grows with it, because aging organs are generally more
prone to failure. However, there is now a critical shortage
of donor organs, and many patients in need of organs will

die while waiting for transplants. In addition, even if an or-
gan becomes available, rejection of organs is still a major
problem in transplant patients despite improvements in
the methods used for immunosuppression following the
transplant procedure. Even if rejection does not occur, the
need for lifelong use of immunosuppressive medications
leads to a number of complications in these patients.
 These problems have led physicians and scientists to
look to new fields for alternatives to organ transplantation.
In the 1960s, a natural evolution occurred in which re-
searchers began to combine new devices and materials sci-
ences with cell biology, and a new field that is now termed
tissue engineering was born. As more scientists from differ-
ent fields came together with the common goal of tissue re-
placement, the field of tissue engineering became more for-
mally established. Tissue engineering is now defined as “an
interdisciplinary field which applies the principles of en-
gineering and life sciences towards the development of bio-


Tissue Engineering

logical substitutes that aim to maintain, restore or improve
tissue function.”1 Then, after the discovery of human stem
cells by Thomson’s group in the early 1980s,2 the field of
stem cell biology took shape and suggested that it may one
day be possible to obtain and use donor stem cells in tissue
engineering strategies, or perhaps even reactivate endoge-
nous stem cells and use them to regenerate failing organs
in adult patients.
 The fields of stem cells, cell transplantation, and tissue
engineering all have one unifying concept-the regeneration
of living tissues and organs. Thus, in 1999, William Haseltine,
then the Scientific Founder and Chief Executive Officer of
Human Genome Sciences, coined the term regenerative
medicine, in effect bringing all these areas under one defin-
ing field.3

 In the past two decades, scientists have attempted to en-
gineer virtually every tissue of the human body. This ar-
ticle will review the basic techniques used in tissue en-
gineering and discuss some of the progress that has been
achieved in this field.


 The field of regenerative medicine encompasses various
areas of technology, such as tissue engineering, stem cells,
and cloning. Tissue engineering, one of the major areas of
regenerative medicine, follows the principles of cell trans-
plantation, materials science, and engineering toward the
development of biological substitutes that can restore and
maintain normal function. Tissue engineering strategies
generally fall into two categories: the use of acellular scaf-
folds, which depend on the body’s natural ability to re-
generate for proper orientation and direction of new tissue
growth, and the use of scaffolds seeded with cells. Acellular
scaffolds are usually prepared by manufacturing artificial
scaffolds or by removing cellular components from tissues
via mechanical and chemical manipulation to produce acel-
lular, collagen-rich matrices.4-7 These matrices tend to
slowly degrade on implantation and are generally replaced
by the extracellular matrix (ECM) proteins that are se-
creted by the in-growing cells. Cells can also be used for
therapy via injection, either with carriers such as hydro-
gels or alone.

1. Biomaterials for use in regenerative medicine
 In the past, synthetic materials were introduced to re-
place or to rebuild diseased tissues or parts in the human
body. The manufacture of new materials, such as tetra-
fluoroethylene (Teflon) and silicone, opened a new field of
research that led to the development of a wide array of de-
vices that could be applied for human use. Although these
devices could provide structural support or replacement,
the functional component of the original tissue was not
restored. However, studies in cell biology, molecular biol-
ogy, and biochemistry allowed a better understanding of
the ECM and its interaction with cells in the tissues of the

body, as well as interactions with growth factors and their
ligands, and as a result, new biomaterials were designed
with these interactions in mind.
 In tissue engineering, biomaterials replicate the bio-
logical and mechanical function of the native ECM found
in tissues in the body. Biomaterials provide a three-dimen-
sional space in which cells can attach, grow, and form new
tissues with appropriate structure and function. They also
allow for the delivery of cells and appropriate bioactive fac-
tors (e.g., cell adhesion peptides, growth factors) to desired
sites in the body.8 Because most mammalian cell types are
anchorage-dependent and will die if no cell-adhesion sub-
strate is available, biomaterials provide this substrate
while allowing delivery of cells with high loading efficiency.
Biomaterials can also provide mechanical support against
in vivo forces so that the predefined three-dimensional
structure of a tissue-engineered organ is maintained dur-
ing tissue development.
 The ideal biomaterial should be biodegradable and bio-
resorbable to support the replacement of normal tissue
without inducing inflammation. Incompatible materials
are destined for an inflammatory or foreign-body response
that eventually leads to rejection or necrosis. Because bio-
materials provide temporary mechanical support while
the cells undergo spatial reorganization into tissue, a prop-
erly chosen biomaterial should allow the engineered tissue
to maintain sufficient mechanical integrity to support it-
self in early development, while in late development, it
should have begun degradation such that it does not hinder
further tissue growth.8 The degradation products, if pro-
duced, should be removed from the body via metabolic path-
ways at an adequate rate to ensure that the concentration
of these degradation products in the tissues remains at a
tolerable level.9

 Generally, three classes of biomaterials have been uti-
lized for engineering tissues: naturally derived materials
(e.g., collagen and alginate),10-14 acellular tissue matrices
(e.g., bladder submucosa and small intestinal submuco-
sa),4-7 and synthetic polymers such as polyglycolic acid
(PGA), polylactic acid (PLA), and poly (lactic-co-glycolic
acid) (PLGA).15-18 These classes of biomaterials have been
tested with respect to their biocompatibility.19,20 Naturally
derived materials and acellular tissue matrices have the
potential advantage of biological recognition. However,
synthetic polymers can be produced reproducibly on a large
scale with controlled properties such as strength, degrada-
tion rate, and microstructure.

2. Cells for use in cell therapy and tissue engineering
 1) Native cells: When native cells are used for tissue en-
gineering, a small piece of donor tissue is dissociated into
individual cells. These cells are expanded in culture and ei-
ther injected directly back into the host or attached to a sup-
port matrix and then reimplanted. The source of donor tis-
sue can be heterologous (such as bovine), allogeneic (same
species, different individual), or autologous. The preferred
cells to use are autologous cells, where a biopsy of tissue


Jennifer L. Olson, et al

is obtained from the host, the cells are dissociated and ex-
panded in culture, and the expanded cells are implanted
into the same host.5,21-28 The use of autologous cells, al-
though it may cause an inflammatory response, avoids re-
jection, and thus the deleterious side effects of immuno-
suppressive medications can be avoided.
 Ideally, both structural and functional tissue replace-
ment will occur with minimal complications when autolo-
gous native cells are used. However, one of the limitations
of applying cell-based regenerative medicine techniques to
organ replacement has been the inherent difficulty of grow-
ing specific cell types in large quantities. Even when some
organs, such as the liver, have a high regenerative capacity
in vivo, cell growth and expansion in vitro may be difficult.
By studying the privileged sites for committed precursor
cells in specific organs, as well as exploring the conditions
that promote differentiation, one may be able to overcome
the obstacles that limit cell expansion in vitro. For exam-
ple, urothelial cells could be grown in the laboratory setting
in the past, but only with limited expansion. Several proto-
cols were developed over the past two decades that identi-
fied the undifferentiated cells and kept them undifferen-
tiated during their growth phase.27,29,30-32 With the use of
these methods of cell culture, it is now possible to expand
a urothelial strain from a single specimen that initially cov-
ered a surface area of 1 cm2 to one covering a surface area
of 4,202 m2 (the equivalent of one football field) within 8
weeks.27 These studies indicated that it should be possible
to collect autologous bladder cells from human patients, ex-
pand them in culture, and return them to the donor in suffi-
cient quantities for reconstructive purposes.27,30 Major ad-
vances have been achieved within the past decade on the
possible expansion of a variety of primary human cells,
with specific techniques that make the use of autologous
cells for clinical application possible.
 Most current strategies for tissue engineering depend on
a sample of autologous cells from the diseased organ of the
host. However, for many patients with extensive end-stage
organ failure, a tissue biopsy may not yield enough normal
cells for expansion and transplantation. In other instances,
primary autologous human cells cannot be expanded from
a particular organ, such as the pancreas. In these sit-
uations, stem cells are envisioned as being an alternative
source of cells from which the desired tissue can be derived.
Stem cells can be derived from discarded human embryos
(human embryonic stem cells), from fetal tissue, or from
adult sources (bone marrow, fat, skin).

3. Stem cells for use in tissue engineering
 1) Embryonic stem cells: Human embryonic stem (hES)
cells exhibit two remarkable properties: the ability to pro-
liferate in an undifferentiated but pluripotent state (self-
renewal), and the ability to differentiate into many speci-
alized cell types.36 They can be isolated by aspirating the
inner cell mass from the embryo during the blastocyst stage
(5 days post-fertilization) and are usually grown on feeder
layers consisting of mouse embryonic fibroblasts or human

feeder cells.37 More recent reports have shown that these
cells can be grown without the use of a feeder layer38 and
thus avoid the exposure of these human cells to mouse vi-
ruses and proteins. These cells have demonstrated lon-
gevity in culture by maintaining their undifferentiated
state for at least 80 passages when grown by use of current
published protocols.2,39 In addition, hES cells are able to dif-
ferentiate into cells from all three embryonic germ layers
in vitro. Skin and neurons have been formed, indicating ec-
todermal differentiation.40-43 Blood, cardiac cells, cartilage,
endothelial cells, and muscle have been formed, indicating
mesodermal differentiation.44-46 Pancreatic cells have been
formed, indicating endodermal differentiation.47 In addi-
tion, as further evidence of their pluripotency, embryonic
stem cells can form embryoid bodies, which are cell ag-
gregations that contain all three embryonic germ layers
while in culture and can form teratomas in vivo.48 However,
there are many ethical and religious concerns associated
with hES cells because embryos are destroyed in order to
obtain them. Thus, the use of these cells is currently banned
in many countries.
 2) Stem cells from somatic cell nuclear transfer: Stem
cells for tissue engineering could also be generated through
cloning procedures. There has been tremendous interest in
the field of nuclear cloning since the birth of the cloned
sheep Dolly in 1997, but actually, Dolly was not the first
animal produced by using nuclear transfer. In fact, frogs
were the first successfully cloned vertebrates derived from
nuclear transfer.49 However, in the frog experiment, the
nuclei used for cloning were derived from non-adult sources.
In fact, live lambs were produced in 1996 by using nuclear
transfer as well, but they were produced from differen-
tiated epithelial cells derived from embryonic discs.50 The
significance of Dolly was that she was the first mammal to
be derived from an adult somatic cell by use of nuclear
transfer.51 Since then, animals from several species have
been grown by using nuclear transfer technology, inclu-
ding cattle,52 goats,53 mice,54 and pigs.55-56

 Two types of nuclear cloning, reproductive cloning and
therapeutic cloning, have been described, and a better un-
derstanding of the differences between the two types may
help to alleviate some of the controversy that surrounds
these technologies.57-58 Banned in most countries for hu-
man applications, reproductive cloning is used to generate
an embryo that has the identical genetic material as its cell
source. This embryo can then be implanted into the uterus
of a female to give rise to an infant that is a clone of the
donor. On the other hand, therapeutic cloning is used to
generate early stage embryos that are explanted in culture
to produce embryonic stem cell lines whose genetic materi-
al is identical to that of its source. These autologous stem
cells have the potential to become almost any type of cell
in the adult body, and thus would be useful in tissue and
organ replacement applications.59 Therefore, therapeutic
cloning, which has also been called somatic cell nuclear
transfer, may provide an alternative source of transplan-
table cells. According to data from the Centers for Disease

Tissue Engineering

Control and Prevention, an estimated 3,000 Americans die
every day of diseases that could have been treated with
stem cell–derived tissues.60-61 With current allogeneic tis-
sue transplantation protocols, rejection is a frequent com-
plication because of immunologic incompatibility, and im-
munosuppressive drugs are usually required.59 The use of
transplantable tissue and organs derived from therapeutic
cloning could lead to the avoidance of immune responses
that typically are associated with transplantation of non-
autologous tissues.60

 While promising, somatic cell nuclear transfer technol-
ogy has certain limitations that require further study be-
fore this technique can be applied widely in tissue or organ
replacement therapy. First, the efficiency of the cloning
process is very low, as evidenced by the fact that most em-
bryos derived from the cloning process do not survive.62-64
To improve cloning efficiency, further improvements are
required in many of the complex steps of nuclear transfer,
such as the enucleation process for oocytes, the actual
transfer of a nucleus to this enucleated oocyte, and the acti-
vation process that instructs the cloned oocytes to begin
dividing. In addition, cell cycle synchronization between
donor cells and recipient oocytes must be accomplished.65

 3) Reprogramming and generation of iPS cells: Within
the past few years, exciting reports of the successful trans-
formation of adult somatic cells into pluripotent stem cells
through genetic “reprogramming” have been published.
Reprogramming is a technique that involves de-differen-
tiation of adult somatic cells (such as fibroblasts) to produce
patient-specific pluripotent stem cells. This process is es-
pecially exciting because it allows pluripotent stem cells to
be obtained without the use of embryos. Also, cells genera-
ted by reprogramming are genetically identical to the so-
matic cells used (and thus to the patient who donated these
cells) and should not be rejected. Yamanaka was the first
to discover that mouse embryonic fibroblasts (MEFs) and
adult mouse fibroblasts could be reprogrammed into an
“induced pluripotent state (iPS).”66 They examined 24
genes that were thought to be important for embryonic
stem cells and identified 4 key genes that, when introduced
into the reporter fibroblasts via retroviral vectors, resulted
in drug-resistant cells. These were Oct3/4, Sox2, c-Myc,
and Klf4. The resultant iPS cells possessed the immortal
growth characteristics of self-renewing embryonic stem
cells, expressed genes specific for embryonic stem cells, and
generated embryoid bodies in vitro and teratomas in vivo.
When iPS cells were injected into mouse blastocysts, they
contributed to a variety of cell types. However, although
iPS cells selected in this way were pluripotent, they were
not identical to embryonic stem cells. Unlike embryonic
stem cells, chimeras made from iPS cells did not result in
full-term pregnancies. Gene expression profiles of the iPS
cells showed that they possessed a distinct gene expression
signature that was different from that of embryonic stem
cells. In addition, the epigenetic state of the iPS cells was
somewhere between that found in somatic cells and that
found in embryonic stem cells, suggesting that the re-

programming was incomplete.
 These results were improved significantly by Wernig
and Jaenisch in July 2007.67 Fibroblasts were infected with
retroviral vectors and selected for the activation of endoge-
nous Oct4 or Nanog genes. Results from this study showed
that DNA methylation, gene expression profiles, and the
chromatin state of the reprogrammed cells were similar to
those of embryonic stem cells. Teratomas induced by these
cells contained differentiated cell types representing all
three embryonic germ layers. Most importantly, the re-
programmed cells from this experiment could form viable
chimeras and contribute to the germline-like embryonic
stem cells, suggesting that these iPS cells were completely
reprogrammed. Wernig et al observed that the number of
reprogrammed colonies increased when drug selection was
initiated later (day 20 rather than day 3 post-transduc-
tion). This suggests that reprogramming is a slow and grad-
ual process and may explain why previous attempts re-
sulted in incomplete reprogramming.
 It has recently been shown that reprogramming of hu-
man cells is possible.68-69 Yamanaka generated human iPS
cells that are similar to hES cells in terms of morphology,
proliferation, gene expression, surface markers, and ter-
atoma formation. Thompson’s group showed that retro-
viral transduction of the stem cell markers OCT4, SOX2,
NANOG, and LIN28 could generate pluripotent stem cells.
However, in both studies, the human iPS cells were similar
but not identical to hES cells. Although reprogramming is
an exciting phenomenon, our limited understanding of the
mechanism underlying it currently limits the clinical ap-
plicability of the technique, but the future potential of re-
programming is quite exciting.
 4) Amniotic fluid and placental stem cells: An alternate
source of stem cells is the amniotic fluid and placenta.
Amniotic fluid and the placenta are known to contain mul-
tiple partially differentiated cell types derived from the de-
veloping fetus. We isolated stem cell populations from these
sources, called amniotic fluid and placental stem cells
(AFPSC), that express embryonic and adult stem cell mar-
kers.70 The undifferentiated stem cells expand extensively
without feeders and double every 36 hours. Unlike hES
cells, the AFPSC do not form tumors in vivo. Lines main-
tained for over 250 population doublings retained long telo-
meres and a normal karyotype. AFS cells are broadly multi-
potent. Clonal human lines verified by retroviral marking
can be induced to differentiate into cell types representing
each embryonic germ layer, including cells of adipogenic,
osteogenic, myogenic, endothelial, neuronal, and hepatic
lineages. In this respect, they meet a commonly accepted
criterion for pluripotent stem cells, without implying that
they can generate every adult tissue. Examples of differen-
tiated cells derived from AFS cells and displaying speciali-
zed functions include neuronal lineage cells secreting the
neurotransmitter L-glutamate or expressing G-protein-
gated inwardly rectifying potassium (GIRK) channels,
hepatic lineage cells producing urea, and osteogenic line-
age cells forming tissue engineered bone. The cells could

Jennifer L. Olson, et al

be obtained either from amniocentesis or chorionic villous
sampling in the developing fetus, or from the placenta at
the time of birth. The cells could be preserved for self-use
and used without rejection, or they could be banked. A bank
of 100,000 specimens could potentially supply 99% of the
US population with a perfect genetic match for transplan-
tation. Such a bank may be easier to create than with other
cell sources, because there are approximately 4.5 million
births per year in the USA.70

 5) Adult stem cells: Adult stem cells, especially hemato-
poietic stem cells, are the best understood cell type in stem
cell biology.71 The presence of stem cells in the adult was
first discerned by Till and McCulloch, who were investigat-
ing the mechanisms by which the bone marrow could re-
generate after exposure to radiation.72 However, adult stem
cell research remains an area of intense study, because
their potential for therapy may be applicable to a myriad
of degenerative disorders. Within the past decade, adult
stem cell populations have been found in many adult ti-
ssues other than the bone marrow and the gastrointestinal
tract, including the brain,73-74 skin,75 and muscle.76 Many
other types of adult stem cells have been identified in or-
gans all over the body and are thought to serve as the pri-
mary repair entities for their corresponding organs.77 The
discovery of such tissue-specific progenitors has opened up
new avenues for research.
 A notable exception to the tissue-specificity of adult stem
cells is the mesenchymal stem cell (MSC), also known as
the multipotent adult progenitor cell. This cell type is de-
rived from bone marrow stroma.78-79 Such cells can differ-
entiate in vitro into numerous tissue types80-81 and can also
differentiate developmentally if injected into a blastocyst.
Multipotent adult progenitor cells can develop into a variety
of tissues including neuronal,82 adipose,76 muscle,76,83 li-
ver,84-85 lungs,86 spleen,87 and gut tissue,79 but notably not
bone marrow or gonads.
 In addition, stem cells derived from adipose tissue may
also be an autologous and self-renewing cell source. Adi-
pose-derived stem cells (ADSCs) have been shown to diffe-
rentiate into a variety of cell phenotypes, and since they are
easily obtained, they show great promise for future types
of reconstructive surgery based on tissue engineering and
there have been several clinical trials using these cells.
Wilson and Mizuno have both provided excellent, detailed
reviews of these.88-89

 Research into more differentiated types of adult stem
cells has, however, progressed slowly, mainly because in-
vestigators have had great difficulty in maintaining adult
non-mesenchymal stem cells in culture. Some cells, such
as those of the liver, pancreas, and nerve, have very low pro-
liferative capacity in vitro, and the functionality of some
cell types is reduced after the cells are cultivated. Isolation
of cells has also been problematic, because stem cells are
present in extremely low numbers in adult tissue.84,90 While
the clinical utility of adult stem cells is currently limited,
great potential exists for future use of such cells in tissue-
specific regenerative therapies. The advantage of adult

stem cells is that they can be used in autologous therapies,
thus avoiding any complications associated with immune


 The simplest regenerative medicine strategies are those
that are based on the actions of cells, which can be implanted
either alone or within a type of carrier material, such as a
hydrogel. These cell therapies are designed to inject or im-
plant healthy cells to replace populations of cells that are
no longer functioning properly owing to disease or injury.
The cells used in these therapies can be autologous cells de-
rived from a tissue biopsy and expanded in culture, or they
can be stem cells from various sources that can be guided
to differentiate into appropriate cell types by using both en-
dogenous and exogenous biochemical cues.
 For example, one area of intense study in regenerative
medicine is the pancreas, because the ability to replace or
regenerate the insulin-producing cells of this organ could
lead to novel treatments or a cure for diabetes. In a series
of exciting experiments, Zhou et al demonstrated that re-
generation of the insulin-producing cells of the pancreas,
the β-cells, may be possible by using cellular reprogram-
ming techniques91 Using a mouse model, they showed that
in vivo activation of a specific combination of three tran-
scription factors (Ngn3, Pdx1, and Mafa) by use of adeno-
viral vectors led to the reprogramming of adult differ-
entiated pancreatic exocrine cells into cells that closely re-
sembled β-cells. These cells were similar to native β-cells
in size, shape, and ultrastructure, and they expressed genes
that are specific to β-cells as well. Interestingly, these cells
secreted insulin and expressed vascular endothelial growth
factor (VEGF), which allowed them to remodel the local
vasculature in a manner similar to native β-cells. In fact,
these reprogrammed cells were able to partially ameliorate
hyperglycemia in diabetic mice, suggesting that repro-
gramming techniques for treating disease may one day be-
come a reality.
 Degenerative muscle diseases such as Duchenne’s mus-
cular dystrophy have devastating effects on quality of life.
To date, these genetic disorders have no suitable treat-
ment. Early enthusiasm for gene therapy interventions
has been tempered by issues of vector toxicity and inade-
quate gene transfer to target muscle cells in vivo. However,
natural mechanisms of muscle repair have suggested that
cell-based therapy could take advantage of natural homing
mechanisms to direct cells to the proper location.92 Experi-
ments using the mdx mouse model, in which the dystrophin
gene is mutated, indicate that injection of normal muscle
precursors and dermal fibroblasts into skeletal muscle can
lead to increased expression of dystrophin and improved
functional outcomes. However, this treatment option re-
quires further studies before it can be widely applied in the
 Although many of these cell therapies are still in the ex-
perimental stage, some are being translated to the clinic

Tissue Engineering

and clinical trials are being performed. Vesicoureteral re-
flux (VUR; a condition in which urine flows backwards from
the bladder into the ureter and kidney) and stress urinary
incontinence are two urologic conditions that can result
from dysfunction of a specific sphincter muscle. When se-
vere, these conditions are repaired surgically. However,
cell-based therapies for both VUR and incontinence would
be an important alternative to surgical repair of these
conditions. Ideally, such a therapy would be easily ad-
ministered by injection and well tolerated by the patient.
The injectable therapy should be non-antigenic, non-mi-
gratory, volume stable, and safe for human use, and in addi-
tion, it should be able to carry cells and serve as a matrix
in vivo.
 Toward this goal, long-term studies were conducted to
determine the effects of injectable chondrocytes for the
treatment of VUR in vivo.93 Chondrocytes were chosen be-
cause the use of autologous cartilage for the treatment of
VUR in humans would satisfy all of the requirements for
an ideal injectable cell-based therapy. Chondrocytes de-
rived from an ear biopsy can be readily grown and expanded
in culture. Neocartilage formation can be achieved in vitro
and in vivo by using chondrocytes cultured on synthetic bio-
degradable polymers. In the VUR experiments, chon-
drocytes were suspended in an alginate matrix and injected
around the vesicoureteral sphincter. In time, normal carti-
lage replaced the alginate as the alginate slowly degraded.
This system was then adapted for the treatment of VUR in
a porcine model.94 These studies show that chondrocytes
can be easily harvested and combined with alginate in vi-
tro, that the suspension can be easily injected cystoscopi-
cally, and that the elastic cartilage tissue formed can cor-
rect the VUR without any evidence of obstruction.
 Two multicenter clinical trials were conducted by use of
this engineered chondrocyte technology. First, patients
with VUR were treated at 10 centers throughout the United
States. The patients had a similar success rate as with oth-
er injectable substances in terms of cure. Cartilage for-
mation was not noted in patients with treatment failure.
Patients who were cured probably had a biocompatible re-
gion of engineered autologous tissue present.95 Secondly,
patients with urinary incontinence were treated endo-
scopically with injected chondrocytes at three different
medical centers. Phase 1 trials showed an approximate
success rate of 80% at 3 and 12 months postoperatively.96


 Tissue engineering strategies are often referred to as
“growing organs in the laboratory.” In these strategies, dif-
ferentiated cells or stem cells are seeded onto a biomaterial
scaffold and this construct is allowed to mature in vitro in
a bioreactor for a short time before implantation in vivo.
These constructs are designed to replace a malfunctioning
organ in its entirety. In recent years, it has been shown that
hollow organs, such as the urinary bladder, urethra, and
blood vessels, can be successfully engineered in the labo-

ratory, and these successes are described below.
 The urethra can be repaired by using tissue-engineered
grafts in several ways. It has been shown that various bio-
materials without cells, such as PGA and acellular colla-
gen-based matrices from small intestine and bladder, can
be used experimentally (in animal models) for the regene-
ration of urethral tissue.7,97-99 Acellular collagen matrices
derived from bladder submucosa have been used experi-
mentally and clinically. In animal studies, segments of the
urethra were resected and replaced with acellular matrix
grafts in an onlay fashion. Histological examination showed
complete epithelialization and progressive vessel and mu-
scle infiltration, and the animals were able to void through
the neo-urethras.7 These results were confirmed in a clini-
cal study of patients with hypospadias and urethral stric-
ture disease.100 Decellularized cadaveric bladder submu-
cosa was used as an onlay matrix for urethral repair in pa-
tients with stricture disease and hypospadias. Patent,
functional neo-urethras were noted in these patients with
up to a 7-year follow-up. The use of an off-the-shelf matrix
appears to be beneficial for patients with abnormal ure-
thral conditions and obviates the need for obtaining autolo-
gous grafts, thus decreasing operative time and eliminat-
ing donor site morbidity.
 Unfortunately, the above techniques are not applicable
for tubularized urethral repairs. The collagen matrices are
able to replace urethral segments only when used in an on-
lay fashion. However, if a tubularized repair is needed, the
collagen matrices should be seeded with autologous cells
to avoid the risk of stricture formation and poor tissue
development.101 In addition, cell-seeded matrices must be
used if the segment of urethra to be replaced is longer than
about 1 cm.102 Recently, Raya-Rivera and colleagues used
tissue-engineered urethras that had been created from pa-
tients’ own cells for tubularized urethral reconstruction. In
this preliminary study, five boys who had urethral defects
were treated. A tissue biopsy was taken from each patient,
and the muscle and epithelial cells derived from the biopsy
sample were expanded and seeded onto tubularized poly-
glycolic acid:poly(lactide-co-glycolide acid) scaffolds to cre-
ate neo-urethras for implantation. The patients then un-
derwent urethral reconstruction with the engineered
urethras. After surgery, these patients were followed for
up to 6 years. All of the patients experienced an increase
in urinary flow rate, and serial radiographic and endo-
scopic studies showed that they maintained wide urethral
calibers throughout the follow-up period and did not develop
strictures. In addition, urethral biopsies were performed
in these patients, and these revealed that by 3 months post-
surgery, the engineered grafts had developed a normal
appearing tissue architecture consisting of a urothelial layer
surrounded by a muscular layer.103

 In addition, similar techniques have been used to create
tissue-engineered bladder constructs. Urothelial and mu-
scle cells can be expanded in vitro, seeded onto polymer
scaffolds, and allowed to attach and form sheets of cells.104
These principles were applied in several studies in which

Jennifer L. Olson, et al

tissue-engineered bladders were used to repair subtotal
cystectomies in beagle dogs.28,105-106 The first clinical ex-
perience in which engineered bladder tissue for cystoplasty
reconstruction in patients was conducted starting in 1999.
A small pilot study of seven patients was reported, using
a collagen scaffold seeded with cells either with or without
omentum coverage, or a combined PGA-collagen scaffold
seeded with cells and omental coverage. The patients re-
constructed with the engineered bladder tissue created
with the PGA-collagen cell-seeded scaffolds showed in-
creased compliance, decreased end-filling pressures, in-
creased capacities, and longer dry periods.107 Although the
experience is promising in terms of showing that engi-
neered tissues can be implanted safely, it is just a start in
terms of accomplishing the goal of engineering fully func-
tional bladders. Further experimental and clinical work is
being conducted.
 Xenogenic or synthetic materials have been used as re-
placement blood vessels for complex cardiovascular lesions.
However, these materials typically lack growth potential,
and may place the recipient at risk for complications such
as stenosis, thromboembolization, or infection.108 Tissue-
engineered vascular grafts have been constructed by using
autologous cells and biodegradable scaffolds and have been
applied in dog and lamb models.109-112 The key advantage
of using these autografts is that they degrade in vivo and
thus allow the new tissue to form without the long-term
presence of foreign material.108 Translation of these techni-
ques from the laboratory to the clinical setting has begun,
with autologous vascular cells harvested, expanded, and
seeded onto a biodegradable scaffold.113 The resultant auto-
logous construct was used to replace a stenosed pulmonary
artery that had been previously repaired. Seven months af-
ter implantation, no evidence of graft occlusion or aneur-
ysmal changes was noted in the recipient. In addition, an-
other group created tissue-engineered blood vessels by us-
ing the cell-sheet multilayer method and then used these
constructs to successfully create vascular access points for
hemodialysis in 10 patients.114 More recently, the same
group completed a larger study on these engineered vessels
for hemodialysis, which indicated that the 1-month and
6-month patency of the grafts was 78% and 60%, respec-
tively, which meets the approved criteria for a high-risk pa-
tient cohort.115

 Finally, few treatment options are currently available
for patients who suffer from severe congenital tracheal pa-
thology, such as stenosis, atresia, and agenesis, due to the
limited availability of autologous transplantable tissue in
the neonatal period. Tissue engineering in the fetal period
may be a viable alternative for the surgical treatment of
these prenatally diagnosed congenital anomalies, because
cells could be harvested and grown into transplantable ti-
ssue in parallel with the remainder of gestation. Chondro-
cytes from both elastic and hyaline cartilage specimens have
been harvested from fetal lambs, expanded in vitro, and
then dynamically seeded onto biodegradable scaffolds.116
The constructs were then implanted as replacement tra-

cheal tissue in fetal lambs. The resultant tissue-engineered
cartilage was noted to undergo engraftment and epithelial-
ization, while maintaining its structural support and pa-
 Recently, Martin Birchall’s group moved this technology
into a human patient with end-stage airway disease.117 This
group was able to remove the cellular material and MHC
antigens from a human donor trachea and, using a speci-
alized bioreactor, seed this acellular matrix with chon-
drocytes and epithelial cells derived from the patient to re-
ceive the graft. This construct was then used to replace the
patient’s left main bronchus. There were no perioperative
complications, and the left lung ventilated normally as
soon as the graft was placed. At 3 months after surgery, the
patient’s lung function was in the normal range for her age
and sex, and she was able to function normally. Although
longer follow-up and larger study populations are needed,
this report indicates that tissue engineering may be a new
option for patients with airway disease.
 However, whereas there has been exciting progress with
tissue engineering techniques for hollow organs, the develop-
ment of methods to generate larger, solid organs with more
complex histological structure has been much more diffi-
cult. A number of issues must be addressed before fully func-
tional, engineered organs such as liver and kidney can be
prepared in the laboratory. First, these organs contain ex-
tremely complex internal structures made up of numerous
cell types arranged in very specific ways, and simple cell-
seeding techniques may not be sufficient for reconstructing
these structures. In addition, the large size of these organs
dictates that the delivery of oxygen and nutrients to each
part of the organ will be a challenge, unless a method for
engineering a functional vascular network within the or-
gan can be found. However, despite the challenges, there
have been some encouraging results from several studies.
For example, the kidney contains multiple cell types and
a complex functional anatomy that renders it one of the
most difficult to reconstruct,21,118 yet we were able to create
a rudimentary form of this organ that appeared to have at
least the filtration properties of the native kidney.
 We applied the principles of both tissue engineering and
therapeutic cloning in an effort to produce genetically iden-
tical renal tissue in a large animal model, the cow (Bos tau-
rus).119 Bovine skin fibroblasts from adult Holstein steers
were obtained by ear notch, and single donor cells were iso-
lated and microinjected into the perivitelline space of donor
enucleated oocytes (nuclear transfer). The resulting blas-
tocysts were implanted into progestin-synchronized recip-
ients to allow for further in vivo growth. After 12 weeks,
cloned renal cells were harvested and expanded in vitro.
Next, the cloned renal cells were seeded on scaffolds con-
sisting of three collagen-coated cylindrical silastic cathe-
ters. The ends of the three membranes of each scaffold were
connected to catheters that terminated into a collecting
reservoir. This created a renal neo-organ with a mecha-
nism for collecting the excreted urinary fluid. These scaf-
folds with the collecting devices were transplanted sub-

Tissue Engineering

cutaneously into the same steer from which the genetic ma-
terial originated and then retrieved 12 weeks after implan-
 At this time, a yellow urine-like fluid was observed col-
lecting within the reservoir of the device. Chemical analysis
of this fluid, including urea nitrogen and creatinine levels,
electrolyte levels, specific gravity, and glucose concentra-
tion, revealed that the implanted renal cells possessed fil-
tration, reabsorption, and secretory capabilities. Histolo-
gical examination of the retrieved implants revealed exten-
sive vascularization and self-organization of the cells into
glomeruli and tubule-like structures. A clear continuity be-
tween the glomeruli, the tubules, and the silastic catheter
was noted that allowed the passage of urine into the collect-
ing reservoir. These studies demonstrated that cells de-
rived from nuclear transfer can be successfully harvested,
expanded in culture, and transplanted in vivo with the use
of biodegradable scaffolds on which the single suspended
cells can organize into tissue structures that are geneti-
cally identical to those of the host. These studies were the
first demonstration of the use of therapeutic cloning for re-
generation of tissues in vivo. However, the size of this de-
vice was small, and the challenge will be to create a larger
device with functioning vasculature and innervations, so
that it can replace all of the myriad metabolic functions of
the kidney.


 The experiences with urethral, bladder, blood vessel, and
tracheal replacement using tissue engineering provide en-
couragement for future efforts to engineer other organs in
the laboratory. These experiences also cast light on unsolved
problems. For example, innervation of tissues and organs
is important for achieving full functionality. In the canine
engineered bladder experiments, the observation of posi-
tive S-100 staining was consistent with growth of neural
structures into the neo-bladders, and bladder function was
restored soon after implantation.28 Innervation of tissue-
engineered constructs has been observed in other systems
such as the small intestine.120 Not only is successful con-
nection with the nervous system important for the func-
tionality of neo-organs, but evidence suggests that it can
enhance tissue regeneration.121-122 The controlled release
of neurotrophic factors is one potential approach to promote
peripheral nerve regeneration and synapse formation with
engineered tissue.123 Direct electrical stimulation has pro-
ven useful in muscle regeneration124 and may have broader
 An even more fundamental issue for the ultimate success
of laboratory-grown organs, particularly those with com-
plex three-dimensional structure, is the provision of ade-
quate oxygen and the generation of new vasculature. It has
been appreciated for some years that in metabolically ac-
tive tissues, the distance over which oxygen typically must

diffuse from a capillary bed to reach a cell is about 0.1 mm,
but that in clinical grafts, the distance from the edge to the
center of the graft is likely to exceed that by at least 50-
fold.125-126 Therefore, with few exceptions (e.g., cartilage),
oxygen is rate-limiting for the viability of grafted cells, and
thus for organ engineering. Neovascularization, an intri-
cate morphogenetic process that allows the formation of ex-
tensively branched vessels, even in an adult, must occur
rapidly and efficiently for a grafted neo-organ to thrive after
implantation.127 Moreover, special measures may be nece-
ssary to ensure survival of grafted tissue during the initial
period after implantation, until a functional vascular bed
is in place. Currently, three types of strategies have been
devised to solve the oxygen supply problem.
 The first strategy involves the use of mechanical or che-
mical sources of oxygen that can support the construct be-
fore and immediately after implantation, until the neo-
vascularization process is completed and can provide the
neo-organ with sufficient blood circulation. An intra-tissue
perfusion system utilizing an array of micro-needles to de-
liver oxygen and nutrients and eliminate waste enhances
the viability and functionality of thick (1 mm) slices of liver
tissue in vitro and might facilitate in vivo grafting.128 In ad-
dition, the use of oxygen-carrying molecules such as per-
fluorocarbons could promote the function of cells in culture
and of encapsulated cells and organ constructs implanted
into animals.129-130 Our laboratory recently showed that a
PLGA film incorporating an oxygen-generating system
(sodium percarbonate) could prevent the necrosis of ische-
mic tissue over several days in vivo.131 We hope to develop
such novel scaffold materials further to support the surviv-
al of large, complex organ constructs in the initial period
after implantation.
 Second, “prevascularization” strategies aim to generate
neo-organs engineered with a preexisting channel struc-
ture to facilitate the generation of a competent vascular
network.130,132 To accomplish this, endothelial lineage cells
can be pre-seeded into the channels or may be recruited in
vivo by using biochemical signals that are embedded in or
released by the scaffold. However, there is still the question
of how to create channels in a way that will be interpreted
as a natural vascular network by the body. One solution
would be to employ decellularized tissue as the scaffold. A
recent study demonstrated that perfusion of an entire
heart with detergents yields an acellular structure in
which the native vascular channels remain intact.133 We
independently devised perfusion-decellularization tech-
nology using liver tissue and have found that the vascular
tree of the whole organ scaffold remains patent and can be
repopulated with large numbers of endothelial cells.134
Alternatively, several technologies can be used to manu-
facture scaffolds with preformed channels, potentially
with cells incorporated, designed to promote neo-vascula-
rization. For example, laser guided “writing” was used to
pattern endothelial cells and promote their aggregation in-
to tubular vessels.135 Similarly, ink-jet-based bioprinting
of cells and biomaterials by thermal ink jet technology can

Jennifer L. Olson, et al

provide remarkable control of the fine structure of engi-
neered tissues, including the generation of intricate vessel
networks.136 We have used layer-by-layer ink jet printing
to produce three-dimensional constructs containing endo-
thelial cells and showed that these develop functional mi-
crovascularization when implanted in vivo, as assessed by
magnetic resonance imaging.137 Electrospinning of living
cells with biomaterials offers similar potential to fabricate
organ structures with pre-patterned vessels.138 Mathe-
matical modeling of scaffolds designed to contain a preex-
isting arteriovenous loop shows how the provision of an
oxygen source within the scaffold can dynamically support
further neo-vascularization and tissue development.139

 Third, it is well established that growth factors such as
VEGF and FGF can promote vascularization in engineered
tissues.140 Recent efforts have extended this approach by
incorporating additional pro-angiogenic molecules into
scaffolds, such as organ-specific ECM from liver to support
sinusoidal endothelial cells.141 Synthetic biomaterials de-
signed to provide signals normally presented by the ECM
will complement, and may eventually supersede, the use
of the native molecules.142

 Finally, several of the clinical trials involving bioengi-
neered products have been placed on hold because of the
costs involved with the specific technology. With a bioengi-
neered product, costs are usually high because of the bio-
logical nature of the therapies involved, and as with any
therapy, the cost that the medical health care system can
allow for a specific technology is limited. Therefore, the
costs of bioengineered products have to be lowered before
they can have an impact clinically. This is currently being
addressed for multiple tissue-engineered technologies. As
the technologies advance over time, and the volume of the
application is considered, costs will naturally decrease.


 Regenerative medicine efforts are currently underway
experimentally for virtually every type of tissue and organ
within the human body. As regenerative medicine incorpo-
rates the fields of tissue engineering, cell biology, nuclear
transfer, and materials science, personnel who have mas-
tered the techniques of cell harvest, culture, expansion,
transplantation, and polymer design are essential for the
successful application of these technologies to extend hu-
man life. Various tissues are at different stages of develop-
ment, with some already being used clinically, a few in pre-
clinical trials, and some in the discovery stage. Recent prog-
ress suggests that engineered tissues may have an expanded
clinical applicability in the future and may represent a via-
ble therapeutic option for those who would benefit from the
life-extending benefits of tissue replacement or repair.


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The technical essay is a review paper that synthesizes and interprets work
on a particular subject area. Therefore, the format is not as standardized as that
of a research paper. By bringing together the most pertinent findings of
numerous papers from diverse journals, a review paper serves as a valuable
summary of research. In writing your essay, interpret the primary journal article
in a series of paragraphs that build on your discussion, giving particular attention
to the problem or topic posed in your introduction. In addition, relate your
findings to previous observations or experiments from the supplemental
references that you have chosen. Discuss briefly any logical implications of the
journal articles for practical application or future studies. A good review paper
not only synthesizes information; it also provides a critical overview of an
important scientific problem.

After you have finished your first draft of your essay, review the structure
of your manuscript. Are the sections arranged in logical sequence? After you
are satisfied with the structure of your essay manuscript, attend to the details:
the paragraphs, the sentences, and the words. Expect to do several drafts of
your paper before you are satisfied with the final product. Good writing is
generally the product of careful rewriting or revising in which you evaluate your
attempts at organizing and expressing your ideas. In the process you end up
scrutinizing the ideas themselves, as well as your own mastery of the subject.


The text of a biological paper usually contains numerous literature
citations, or references, to the published studies of other authors. This is
because scientists rarely work in a vacuum; hypotheses are developed, tested,
and evaluated in the context of what other scientists have written and discovered.
Thus, careful documentation, or acknowledgment of the work of others, is
essential to good scientific writing. Biologists also need to provide literature
citations because, like other writers, they have an ethical and legal obligation to
give credit to others for material that is not their own. Such material includes not
only direct quotations, but also findings or ideas that stem from the work of
someone else.

Unlike writers in the humanities and social sciences, biologists rarely use
footnotes or endnotes to acknowledge sources. Instead, they insert literature
citations directly in the text, either by giving the last name of the author(s) and the
year of publication (name-and-year method), or by referring to each source by a
number method. Such rules, even if they seem arbitrary, make the reporting of


references an orderly activity, minimizing confusion for writers, readers, editors,
and publishers.

In this course, the name-and-year method, also known as the Harvard
method, will be used for literature citations. Cite each reference by giving the last
name(s) of the author(s) followed by the year in which the material was published.

Work by One Author
For each citation, use parentheses to enclose the name and the date.

Example: The most recent study of sexual dimorphism in this species
(Jackson 1976) fails to account for…

If the author’s name appears as part of the sentence, put just the date in

Example: Black-horned locusts were first reported in Iowa by Blum (1914).

Work by Two Authors
Put the senior authors name first. The senior author is the one whose name
appears first after the title.

Example: In a study by Rutowski and Abrams (1963)

Work by Three or More Authors
Here, you may cite the senior author’s name followed by the abbreviation et al.
(from the Latin phrase et alia, meaning “and others”).

Example: White-lined bark beetles are attracted to the odor of rotting
wood (Bateson et al. 1972)

Whichever documentation system you use, put each citation close to the
information you wish to acknowledge. Do not automatically put cited material at
the end of every sentence. Citations allow you to acknowledge the work or ideas
of others, and they also inform the reader. Do not pack your text with citations
simply to demonstrate that you’ve done your homework and are intimately
familiar with the literature. If certain material is well known and fundamental to a
particular field, it is not necessary to cite sources.



The Literature Cited section of a biological paper contains only the
literature (sources) that have been cited (referred to) in the text of your technical
essay. Even if you have acquired useful background knowledge by reading
several articles and books, do not list any of these in the Literature Cited section
unless you have specifically mentioned them in the text. Bibliographies, lists of
all sources mentioned along with additional references on the topic, are not
generally part of scientific papers.

Biological journals have adopted various formats for the Literature Cited
section of papers. Prospective authors prepare this section by carefully following
the guidelines prescribed by the journal for which they are writing. Your
instructor, like a journal editor, has certain preferences. The following examples
illustrate the style used by the Council of Biological Editors (CBE) Style Manual
(Third Edition).

Journal Article with Single Author

Grand, P. R. 1981. Speciation and the adaptive radiation of Darwin’s
finches. Am. Sci. 69:653-663.

Wolfram, S. 1984. Computer software in science and mathematics. Sci.
Am. 251(3):188-203.

Journal Article with Multiple Authors

Via, S. and R. Lande. 1985. Genotype-environmental interaction.
Evolution 3 :505-522.

Book, Number of Pages Given

Schwarts, R. J. 1955. The complete dictionary of abbreviations. Thomas
Crowel Co., New York. 211 pp.

Book, Part of

Overstreet, H. A. 1925. The psychology of effective writing. Pages 87-
101 in H. A. Overstreet, Influencing human behavior. W. W. Norton & Co.,
New York.


Technical Report

Cowardin, L. M., V. Carter, F. C. Golet, and E. T. LaRoe. 1979 Dec.
Classification of wetlands and deepwater habitats of the United States.
Washington: Fish and Wildlife Service. Report nr FWS/OBS/-79/31. 103


After making final revisions on your paper, you may feel that all the work is
over. You will need to produce a neat, clean manuscript. A sloppily presented
paper, such as one with page numbers missing, margins askew, and the type
barely visible, will not show off your prose to good advantage. Such details affect
the reader’s overall impression of your work.

Use an 8.5-by-11-inch white bond paper (not erasable bond or onionskin),
a dark ink cartridge on your printer, and a standard typeface (not script type or all
capitals). Print on only one side of the paper. Leave two spaces after each
period and a single space after each comma. Indent each paragraph five spaces.
Double-space the entire manuscript, including the abstract (if you have one), the
Literature Cited section (between adjacent references as well as between the
lines of a single reference), table titles, figure legends, and any indented
quotations in the text. Leave margins of 1-1.5 inches on all sides of the page.
Number the pages consecutively beginning with the title page (which does not
actually carry a number, but is still counted). Use only Arabic numerals (1, 2,
3,…) and put the numbers in the upper right-hand corner of each page.

An acceptable format for the technical essay calls for the following
information on the title page: title of paper, author, course information, and date.
Center each line on the page, and capitalize the first letter of all important words.

Typographical errors, misspelled words, missing commas or periods,
irregular spacing, and other minor errors distract the reader and undermine your
authority as a writer. The aim of proofreading is to eliminate such mistakes from
the final draft of the manuscript.

Finally, use a staple to fasten everything together; a separate folder is not

necessary. In case your instructor misplaces the paper, keep a copy of the
original manuscript.

JWF 2012

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